Detector systems for radiation imaging

ABSTRACT

Detector designs and systems for enhanced radiographic imaging with integrated detector systems incorporate one or more of Compton and nuclear medicine imaging, PET imaging and x-ray CT imaging capabilities. Detector designs employ one or more layers of detector modules comprised of edge-on or face-on detectors or a combination of edge-on and face-on detectors which may employ gas, scintillator, semiconductor, low temperature (such as Ge and superconductor) and structured detectors. Detectors may implement tracking capabilities and may operate in a non-coincidence or coincidence detection mode.

CROSS-REFERENCE TO RELATED APPLICATIONS

This application is a continuation-in-part of U.S. patent application Ser. No. 13/573,981, COMPTON CAMERA DETECTOR SYSTEMS FOR NOVEL INTEGRATED COMPTON-PET AND CT-COMPTON-PET RADIATION IMAGING, filed Oct. 18, 2012, which claims priority to U.S. Provisional Application No. 61/689,139, filed May 31, 2012, and U.S. Provisional Application No. 61/690,348, filed Jun. 25, 2012, each of which is incorporated by reference herein, in the entirety and for all purposes. This application is related to co-pending U.S. patent application Ser. No. ______ entitled DETECTOR SYSTEMS FOR RADIATION IMAGING, attorney docket number P247584.US.03, filed on even date herewith, and incorporated by reference herein.

FIELD

This invention provides novel implementations of detector systems that can be employed for diagnostic medical and small animal imaging (diagnostic x-ray radiology including x-ray area, slit, slot, tomosynthesis, CT, phase imaging, intraoral/extraoral dental and radiation therapy imaging, nuclear medicine imaging, PET imaging, small animal imaging), as well as in industrial, Homeland Security and scientific radiation imaging.

BACKGROUND

The combining of imaging modalities to offer increased functionality has produced a number of useful imaging systems, particularly in medical diagnostic and small animal imaging. For example, Gamma ray PET detector systems are frequently sold with x-ray computed tomography (CT) detector systems (although the PET and CT detector systems are physically separate and therefore do not share detectors or a common imaging space). A notable attempt at offering an integrated imaging system (in which detectors and the imaging space of the system are shared) was a SPECT-PET (nuclear medicine and PET) imaging system which reduced costs by sharing detectors and the imaging space (the volume in which the object is imaged). Although these SPECT-PET imaging systems were not well received commercially due to performance compromises nonetheless they offered interesting functionality since SPECT and PET images could be acquired separately or simultaneously in a shared imaging space (thereby avoiding registration error between separately acquired SPECT and PET images and reducing the total scan time). In addition, simultaneous CT-SPECT systems have been proposed (typically using CZT or CdTe) although issues arise due to generally differing collimation and flux rate requirements.

SUMMARY

The invention utilizes improvements in high speed detector electronics along with detector materials developed for human and small animal medical diagnostic imaging including diagnostic x-ray radiology (such as x-ray area, slit, slot, tomosynthesis, CT, dental and phase imaging), radiation therapy imaging, nuclear medicine imaging and PET imaging as well as high energy physics, inspection, etc. to develop cost-effective, single purpose and multipurpose integrated detector systems.

The details of one or more embodiments of the invention are set forth in the accompa-nying drawings and the description below. Other features, objects, and advantages of the invention will be apparent from the description and drawings, and from the claims.

All publications, patents and patent applications cited herein are hereby expressly incorporated by reference for all purposes.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1 illustrates a perspective view of a non-coincidence Compton-PET detector imaging system.

FIG. 2 illustrates a perspective view of an edge-on silicon detector substrate in which shielded readout ASICs are mounted within an etched region along the bottom edge of the semiconductor detector substrate.

FIG. 3 illustrates a perspective view of a focused planar detector.

FIG. 4 illustrates a perspective view of a coincidence Compton-PET detector imaging system.

FIG. 5 illustrates a perspective view of a non-coincidence CT-Compton-PET detector imaging system.

FIG. 6 illustrates a perspective view of a minifying scintillating fiber array coupled to a 1D photodetector structured detector suitable for PC or limited PCE CT imaging.

FIG. 7 illustrates a perspective view of a one-dimensional structured mold detector system with quantum dots or semiconductor detector materials.

FIG. 8 illustrates a perspective view of a two-dimensional structured mold detector system with quantum dots or semiconductor detector materials.

FIG. 9 illustrates a perspective view of a multilayer detector system with N=4 layers used for CT and/or PET detector imaging.

FIG. 10A illustrates a perspective view of an alternate multilayer detector system with N=3 layers used for CT and/or PET detector imaging.

FIG. 10B illustrates a perspective view of a multilayer CT and/or PET detector imaging system with a face-on back-end detector layer.

FIG. 10C illustrates a perspective view of a multilayer CT and/or PET detector imaging system with a face-on back-end detector layer.

FIG. 10D illustrates a perspective view of a multilayer CT and/or PET detector imaging system with a face-on back-end detector layer.

FIG. 11 illustrates a perspective view of a focused two-dimensional structured mold detector system with quantum dots or semiconductor detector materials.

DETAILED DESCRIPTION

Compton cameras are frequently implemented as multilayer detectors. Photon-tracking Compton camera designs considered for photon energies encountered in applications such as nuclear medicine and PET imaging include a single layer (a front-end detector) which provides 3D detector properties by incorporating a stack of face-on detector planes of the same material such as low-Z Silicon (Si) or moderate-Z Germanium (Ge), essentially a multilayer detector, and a multilayer (dual-layer) configuration which combines a 2D detector first layer (the front-end detector) and a 2D detector second layer (the back-end detector). Note that the spatial resolution of the first layer and the second layer in the multilayer (dual-layer) detector design need not be the same. Furthermore, spatial resolution of detector elements within a layer need not be the same (e.g., a detector layer that offers higher spatial resolution in the center or a detector layer in which the pixel size increase with depth).

The dual-layer, front-end/back-end detector configuration typically consists of a face-on, planar, 2D Si (low-Z) or 2D Ge (moderate-Z) front-end detector combined with a face-on, 2D high-Z back-end detector. Thus, these two Compton camera configurations described herein can utilize detector layers of the same material (low-Z and moderate-Z for Compton scattering) or different materials (low-Z for Compton scattering and high-Z for photoelectric interactions) for the detection of photons in the diagnostic energy range of medical imaging. Flexibility in the selection of detector materials and configuration (often with different temporal and/or energy resolution) is not limited to separate layers, and different detector materials and configurations can be employed within a detector layer.

Clearly other choices of materials can be made depending on the photon energy range or if other types of particles (neutrons, muons, etc.) are to be detected. Compton camera designs (as well as x-ray scanning and CT, SPECT, PET and hand-held probe designs are described in various U.S. patents and patent applications including: R. S. Nelson and Z. L. Barbaric, U.S. Pat. No. 4,560,882; R. S. Nelson, U.S. Pat. No. 4,937,453; R. S. Nelson, U.S. Pat. No. 5,258,145; R. S. and W. B. Nelson, U.S. Pat. No. 6,583,420; R. S. and W. B. Nelson, U.S. Pat. No. 7,291,841; R. S. Nelson, U.S. Pat. No. 7,635,848; R. S. and W. B. Nelson, U.S. Pat. No. 8,017,906; R. S. Nelson, U.S. Pat. No. 8,115,174; R. S. Nelson, U.S. Pat. No. 8,115,175; R. S. Nelson, U.S. Pat. No. 8,183,533; R. S. and W. B. Nelson, U.S. patent application Ser. No. 13/199,612, filed Sep. 6, 2011 (U.S. Publication No. 2012/0181437); and R. S. and W. B. Nelson, U.S. patent application Ser. No. 13/507,659, filed Jul. 18, 2012 (U.S. Publication No. 2013/0028379), which are incorporated by reference herein.

Compton camera detector systems exploit the Compton scatter interaction and can also exploit photoelectric interactions (and even pair production interactions at sufficiently high photon energies). Compton camera detector systems include the capability to track these interactions in terms of spatial location and energy deposition with a temporal resolution limited by the detector itself and the readout electronics.

Typically the interaction information is used to estimate the directionality and energy of the photon incident on the Compton camera detector system whether the photon is an x-ray, a gamma ray, or an annihilation gamma ray. Note that with the addition of collimation such as (for example) a pin hole or parallel hole collimator, the Compton camera can be converted into a nuclear medicine SPECT camera (Gamma camera). Compton camera features such as tracking capability can continue to be utilized. This is an example of a dual-use, integrated Compton detector system in which the types of applications are relatively unchanged but the capabilities of the detector system are modified (Nelson, U.S. Pat. No. 7,291,841; U.S. Pat. No. 8,017,906).

The collimation now provides the directionality of an incident gamma ray independent of directionality determined by applying Compton camera reconstruction algorithms. It will be shown that the integrated Compton detector system design can be applied to a range of applications (including nuclear medicine). By employing two or more Compton camera detector systems with electronic coincidence circuitry (used in medical PET detector systems) coincidence PET imaging can be implemented.

The flexibility of the Compton camera detector system design allows versatile non-coincidence Compton-PET and coincidence Compton-PET detector systems to be implemented. Furthermore, CT capability can be implemented in the Compton camera detector system design, including non-coincidence and coincidence Compton-PET designs resulting in CT-Compton-PET detector systems. A simplification of this design in which the CT detector and the Compton-PET detector (or just a PET detector) function independently will be referred to as a limited CT-Compton-PET detector system. Furthermore, limited implementations of Compton camera detector designs can be employed for dedicated applications such as (but not limited to) CT imaging or PET imaging.

Although applications discussed herein are primarily directed at medical diagnostic x-ray and gamma ray radiation detection, in principle the invention can also be used to detect radiation such as charged particles (alphas, betas, protons, muons, etc.) and neutrons (as well as other neutral particles) for the applications described. Furthermore, the detector systems described herein can be combined with or integrated with other imaging modalities such as MRI scanners, optical scanners, ultrasound scanners, opto-acoustic scanners, microwave scanners, etc. It should be understood that the variations of the dual-use detector systems (triple-use, etc. detector systems can also implemented) described herein can be employed for simultaneous or non-simultaneous imaging as required by the appropriate application.

The invention provides detector designs that employ one or more layers of detector modules comprised of edge-on or face-on (or tilted) detectors or a combination of edge-on and face-on detectors (as well as tilted detectors). Edge-on detectors (and tilted detectors) can incorporate sub-aperture resolution (SAR) capabilities and face-on detectors can incorporate depth-of-interaction (DOI) capabilities. One or more types of detectors can be employed, including: scintillator detectors, semiconductor detectors, gas detectors, low temperature (such as Ge or superconductor) detectors and structured detectors.

Detectors can offer block, 1D, 2D or 3D spatial resolution as well as adequate, fast or very fast temporal resolution (depending on the application requirements). Detectors can offer fixed or adjustable pixels sizes which can be uniform or non-uniform (for example, increasing pixel length along the depth dimension as a function of depth to compensate for beam hardening with depth in a CT detector). The effective pixel length along a detector column can be synthesized from the outputs of one or more uniformly spaced pixels. Parallel or focused pixel structures can be implemented. Detectors can operate as energy integrators, photon counters (PCs) and photon counters with energy resolution (PCEs). Possible detector formats include, but are not limited to, planar (and focused planar) and focused structure (parallel planes, ring, partial ring as well as focused ring and focused partial ring) detector geometries.

The invention provides novel detector designs and systems for enhanced radiation imaging including Compton and nuclear medicine imaging, PET imaging and x-ray CT imaging. The invention also provides integrated detector systems based on Compton camera designs.

In one aspect, the invention provides integrated non-coincidence Compton-PET detector imaging systems. In another aspect, the invention provides integrated coincidence Compton-PET detector imaging systems. In yet another aspect, the invention provides limited integrated CT-Compton-PET detector imaging systems. In still another aspect, the invention provides integrated non-coincidence CT-Compton-PET detector imaging systems. In another aspect, the invention provides integrated coincidence CT-Compton-PET detector imaging systems. Since the integrated nature of these Compton camera detector design implementations is readily apparent the term “integrated” will frequently be omitted when referring to them. Therefore “integrated non-coincidence Compton-PET detector imaging systems” will also be referred to as “non-coincidence Compton-PET detector imaging systems,” etc. In still another aspect, the invention provides variations of Compton camera detector designs that can be implemented for dedicated applications such as (but not limited to) CT imaging or PET imaging.

The invention employs a range of detector types and formats. The use of gas, scintillator, semiconductor, low temperature (such as Ge and superconductor) and structured detectors in edge-on and/or face-on geometries has been described for both medical and non-medical imaging applications. Medical imaging applications include diagnostic x-ray imaging (such as slit scanning, slot scanning, area radiography, flat panel or planar cone beam CT, focused structure ring or partial ring fan beam CT, cone beam CT, tomosynthesis, phase, radiation therapy and intraoral/extraoral dental imaging), nuclear medicine imaging (Compton camera, SPECT/gamma camera detector imaging systems as well as hand held probe detectors) and PET imaging. Non-medical imaging applications include high energy physics, x-ray and gamma ray astronomy, industrial radiography, Home Land Security (HLS) and military applications. Furthermore it has been shown that detector spatial resolution can be enhanced using sub-aperture resolution (SAR) or depth-of-interaction (DOI) readout techniques with edge-on and face-on detector geometries, respectively.

Detectors may be layered (stacked) and detector modules within a layer can be partially or completely offset from neighboring detector modules. Individual detectors may function as energy integrators, photon counters (PCs) or photon counters with energy resolution (PCEs), depending on the application. One or more of these detector types can be employed within a detector imaging system. (Photon counting (PC) is often mixed up with photon counting with energy resolution (PCE) in the literature. PC functions as a (one energy bin) single channel analyzer (SCA). PCE functions as a multi-channel analyzer (MCA)).

High speed electronics is provided for tracking interactions and analyzing the readout signals. An electronic communications link is provided to a computer for data post-processing, storage, and display. One or more tracking capabilities such as examining nearest neighbor pixels for effects related to induced signals and charge diffusion, following scattered or characteristic x-ray radiation within a detector layer and between detector layers (if there is more than one detector layer), following Compton scattered electrons and photoelectrons and measuring coincidence events (for example, the detection of pairs of annihilations photons in PET imaging), etc., can be implemented. Tracking techniques are used in photon counting and spectral x-ray imaging, SPECT, PET, Compton cameras, hand-held radiation detector probes, neutron detectors, detectors with SAR or DOI capability and high energy physics particle detectors.

Various Compton camera implementations incorporate one or multiple detector layers. These detector layers provide suitable 2D or 3D spatial resolution, energy resolution, temporal resolution, stopping and scattering power and tracking capability. Compton camera, nuclear medicine SPECT/gamma camera and PET detector imaging systems, tracking, x-ray CT and slit and slot scan detectors, hand held probe detectors, edge-on and face-on detectors (with or without SAR or DOI capability), integrating, PC, and PCE detectors, multi-material detectors along with planar and focused structure detector geometries have been described in various U.S. patents and patent applications including Nelson et al., U.S. Pat. No. 4,560,882; U.S. Pat. No. 4,937,453; U.S. Pat. No. 5,258,145; U.S. Pat. No. 6,583,420; U.S. Pat. No. 7,291,841; U.S. Pat. No. 7,635,848; U.S. Pat. No. 8,017,906; U.S. Pat. No. 8,115,174; U.S. Pat. No. 8,115,175; U.S. Pat. No. 8,183,533; U.S. patent application Ser. No. 13/199,612 and U.S. patent application Ser. No. 13/507,659, which are incorporated by reference herein.

X-ray or gamma ray interactions (in medical imaging applications) can be tracked between sufficiently thin detector layers, each with 2D spatial resolution capability. If the depth of a 2D detector layer is sufficiently small such that tracking position errors are acceptable then it effectively functions as a restricted 3D detector (its depth resolution is at most the thickness of the detector layer). If detectors offer 3D spatial resolution capability then interaction tracking (including multiple interactions) can be implemented internally within a 3D detector layer as well as between detector layers.

Energy resolution can be used to measure the position-dependent energy losses due to the interactions within detectors which in turn can provide an estimate of the energy of the initial incident x-ray or gamma ray. This information can be used to determine whether the initial incident photon energy is within an allowed energy range as well as its directionality.

Temporal resolution capability can be used to distinguish between independent incident x-rays or gamma rays interactions (as well as their subsequent interactions) within the Compton camera. It will be shown that very good temporal resolution can be beneficial if coincidence timing is of interest between Compton camera systems (for example, when coincidence PET imaging is implemented).

One implementation of a Compton camera using a dual-layer detector design wherein the first layer (front end) was a small area, face-on, Si or Ge semiconductor pixelated detector offered 2D spatial resolution. The second layer (back end) was a large area, face-on, scintillator (gamma camera) detector which also offered 2D spatial resolution (Singh, M., Medical Physics Vol. 10(4), pp. 421-427 (July/August 1983) and Singh, M., Doria D., Medical Physics Vol. 10(4), pp. 428-435 (July/August 1983)). Both front-end and back-end detectors offered appropriate levels of energy resolution for the photon energies employed and temporal resolution for the expected event interaction rates.

Since Compton scattered photons include a range of scatter angles the sensitivity of design is in part dependent on the separation distance and area of the second layer with respect to the first layer of detectors. A second layer which employs a smaller 3D detector may, in some instances, be more-cost effective than a larger 2D detector which suffers from parallax errors and needs to be positioned further away from the first layer.

Another implementation of the Compton camera, the (face-on) Compton telescope camera, consisted of only a first layer detector. This front-end detector was comprised of a stack (and thus could also be viewed as a multilayer detector) of 2D, face-on Si detectors which function together as a 3D detector (Kroeger R, et al., IEEE Trans. Nucl. Sci., Vol. 49(4), pp. 1887-1892 (August 2002); Nelson, U.S. Pat. No. 8,017,906).

A stack of 2D, face-on Ge detectors (or a thick 3D Ge detector with DOI capability) can also be implemented although the Ge semiconductor may operate at a low temperature. The Compton telescope camera tracks multiple Compton scatters by a photon in order to determine its original direction and energy.

Low-Z (atomic number) semiconductor materials such as Si and diamond (and sometimes moderate-Z Ge) are often preferred for the front-end Compton scatter detector for photons of relatively low energies (e.g. medical diagnostic x-ray energies, 140.5 keV gamma rays from Tc-99m used in nuclear medicine) compared to the 511 keV gamma rays used in PET imaging.

The Compton scatter interaction cross section of the material dominates its photoelectric cross section and the relative contribution to the angular reconstruction error due to the Doppler shift is reduced as Z decreases and/or photon energy increases. As the photon energy increases semiconductor materials with moderate-Z values (such as Ge, GaAs, CdTe, CZT, etc.) represent increasingly acceptable substitutes for low-Z semiconductor materials such as Silicon.

The amount of energy deposited by relatively low energy photons (commonly used in diagnostic x-ray imaging or nuclear medicine) due to a Compton scatter interaction is typically small and therefore semiconductors detectors are employed as front-end detectors because of their superior energy resolution compared to most scintillator detectors. In the dual-layer Compton camera design lower-cost 2D scintillator detectors may be employed in place of semiconductor detectors as back-end detectors if they offer suitable spatial, temporal and energy resolution and stopping power.

The semiconductor front-end detector may be replaced by a low temperature front-end detector or by a scintillator (or gas) front-end detector although energy resolution may suffer. Any significant reduction in accuracy of the calculated incident photon directionality by Compton reconstruction algorithms can be augmented or supplanted by additional information such as coincidence between detectors (used in coincidence PET imaging).

Compton electron tracking in a gas detector can be implemented although this is typically very time-consuming. Cherenkov radiation, despite the relatively weak optical signals, can be exploited for time-of-flight (TOF) measurements. (Cherenkov radiation can be detected when generated in optically-transparent mediums including fluids such as liquids and gases, scintillators and non-scintillators such as transparent plastics, glasses, fibers, diamond films, etc. Thus, transparent dielectric mediums other than scintillators and gases can be also be employed as Compton scatter or photoelectron detectors within a Compton camera detector system although energy resolution could suffer based on the detection of Cherenkov radiation alone. Inexpensive dielectric materials may be acceptable for those applications in which radiation scatter within the object is of reduced importance and therefor lower detector energy resolution is acceptable. Variations of detector designs described herein can include measuring only a Cherenkov signal or a Cherenkov signal and a fluorescence signal or an electronic signal.)

Potential advantages of this dual-layer implementation of a multilayer design may include a less-expensive front-end detector and/or a front-end detector that offers a feature such as fast (greater than 1 nanosecond) or very fast (less than 1 nanosecond) temporal resolution. Very fast temporal resolution is of interest for TOF PET. Benefits of TOF PET include improved image resolution and lower patient radiation dose. Furthermore, the use of coincidence information can also simplify the requirements of the back-end detector.

Compton electron tracking can also be implemented within a detector layer and between detector layers that employ at least one of scintillator-photodetector detector, semiconductor, structured and low temperature detectors. Since electrons readily interact with matter electron tracking is preferably implemented when detecting energetic photons which are Compton scattered, typically generating more-energetic electrons with a more-directional nature. (A similar concept applies to energetic photoelectric interactions which typically generate more-energetic photoelectrons with a more-directional nature. Thus, a Compton camera could utilize sufficiently energetic photoelectric interactions for image reconstruction by tracking the highly directional photoelectrons.)

The tracking of Compton scattered electrons as well as Compton scattered photons can be simplified by enabling longer path lengths for the scattered particles, improving the estimates of scattering angles. Examples of relatively thin, edge-on detector configurations that incorporate gaps between adjacent detectors (including partially- or completely-offset detectors) are shown in FIGS. 1, 3, 5.

Face-on detector configurations with gaps between detector layers can also be implemented. Compton camera image reconstruction can be improved if both the Compton scattered photon and electron are tracked since the solution can be limited to a fraction of a cone surface rather than the full cone surface.

The flexibility of the Compton camera design can be understood by considering front-end (single layer) detector and front-end with back-end (dual-layer) detector implementations of multilayer, edge-on detector Compton camera designs which can be used for low energy and high energy photon imaging. In one dual-layer implementation the front-end detector is used to detect low energy x-rays or gammas and the back-end detector acts to detect higher energy gammas as an edge-on SPECT/gamma camera or PET camera (Nelson, U.S. Pat. No. 7,291,841).

Front-end detectors and back-end detectors can be differentiated based on functionality and/or position. The front-end and back-end detectors should have at least one different property such as position, size, geometry (planar, box, partial-box, ring, partial-ring, etc.), directionality (focused, unfocused), spatial resolution, temporal resolution, energy resolution, interaction probability (material density, thickness, interaction coefficients), orientation (edge-on, face-on, tilt), noise characteristics, detector operation (integrator, PC, PCE), etc.

A multilayer detector can include one or more front-end detectors and back-end detectors. Detector properties within a detector layer can be uniform or non-uniform (continuous, discontinuous, mixing one or more of detector materials, detector operation capabilities, detector orientations, detector temporal characteristics, etc.). A special case of a multilayer detector is a single detector layer that incorporates one or more front-end detectors and back-end detectors. This can be implemented in structured detectors (such as edge-on structured semiconductor detectors including structured 3D semiconductor detectors and structured mold semiconductor detectors, structured scintillator detectors including 3D edge-on or face-on stacked cross-coupled scintillator rod detectors, multilayer scintillator block detectors, scintillating fiber bundle detectors, straw array detectors, etc.).

For example, stacked cross-coupled scintillator rods can vary the scintillator rod properties (material, interaction probabilities, density, temporal characteristics, brightness, etc.) as a function of depth (as well as within a layer and even within individual rods). Front-end stacked cross-coupled layers might use, for example, scintillator(s) preferred for lower energies encountered in SPECT or very fast scintillator(s) suitable for TOF PET while back-end stacked cross-coupled layers might use scintillator(s) preferred for moderate or fast or very fast PET. Furthermore, scintillator rod properties can be varied within at least one of a rod, a layer, between cross-coupled layers.

Varying scintillator temporal characteristic as a function of position could be used to improve event localization based on both optical signal sharing and different temporal decay characteristics of scintillator rods. Consider a planar or ring multilayer detector geometry with two (discontinuous) detector layers in which adjacent 3D edge-on silicon detector modules with PCE capability in the front-end and back-end detector layers are tilted with respect to one another to achieve a focused detector geometry with respect to diverging radiation from at least one source, with the adjacent detector modules in the back-end detector layer offset to fill gaps between the adjacent modules in the front-end detector layer and define a substantially continuous detector configuration. (Optionally, these two layers can be treated as a single detector layer.) Furthermore, consider a multilayer detector with three detector layers (treat the two focused 3D edge-on silicon with PCE capability detector layers as a single detector layer followed by a 2D face-on scintillator with integration capability followed by a 3D edge-on scintillator with PCE capability), employed as a PET camera and x-ray CT imaging system. The 3D edge-on silicon layer and 2D face-on scintillator layer both function as the front-end detector for CT (experiencing different energy spectrums) and alternatively one or both layers could be employed in a dedicated CT imaging system. The 3D edge-on silicon layer also functions as a front-end detector for PET (detecting gammas or scattering gammas). The 3D edge-on scintillator layer acts as the back-end detector for PET (detecting unscattered gamma rays and scattered gamma rays due to the 3D edge-on silicon layer).

A focused, edge-on Compton camera design is described that can employ one or multiple (of the same or different materials) detector layers as well as implementing additional features such as the offset (complete or partial) of adjacent (neighboring) detector modules within a layer. Completely offset detector modules can be used to create two or more detector layers (offset layers) which when employed together can approximate a continuous detector, and therefor can be referred to as either a single layer or two layers (front-end and back-end layers) of detector modules. The offset layer feature of an edge-on Compton camera design can be implemented in PC, PCE and energy integration versions of diagnostic CT detector, including ring and cone beam CT as well as tomosynthesis, PET, CT-PET, Compton-PET, Compton-PET-CT, gamma camera, etc. (e.g., as described in Nelson, U.S. Pat. No. 7,291,841; U.S. Pat. No. 7,635,848; U.S. Pat. No. 8,017,906; U.S. Pat. No. 8,115,174; U.S. Pat. No. 8,115,175; and U.S. Pat. No. 8,183,533; see also, e.g., Danielsson, U.S. Publication No. 2010/0204942 and Bornefalk, U.S. Publication No. 2010/0215230). This complete or partial offset feature can be used for not only edge-on detector implementations but also face-on detector implementations for ring and cone beam CT (for example, a planar or cylindrical arrangement of linear arrays of face-on detectors, each oriented parallel to the axial direction of the scanner) as well as tomosynthesis.

Implementations of the Compton camera design are described herein that are suitable for use as Compton-PET imaging systems and CT-Compton-PET imaging systems. In addition, the positioning of nuclear medicine collimator hardware such as focused, parallel or pin hole collimators between the object being imaged and the Compton camera permits the system of collimator and Compton camera to provide nuclear medicine imaging capabilities (the imaging capabilities of a SPECT/Gamma camera) for those applications in which the Compton camera does not offer adequate imaging properties.

Limited implementations of the Compton camera designs described herein include versions that function only as CT or PET (and SPECT) detector designs. The Compton camera imaging systems described herein will find use in diagnostic medical x-ray CT, nuclear medicine and PET imaging, x-ray micro-CT imaging, dental CT, medium and small animal imaging, radiation therapy imaging, industrial imaging, HLS and military imaging, and scientific imaging.

Compton-Pet Detector Systems

One implementation of the Compton camera is referred to as the Compton-PET detector system (Nelson, U.S. Pat. No. 7,291,841). The Compton-PET detector system design allows flexibility in the choice of detector materials as well as detector geometries. This flexibility is constrained by the intended imaging applications (such as PET only, nuclear medicine and PET, x-ray and PET).

Face-on, edge-on, and combinations of face-on and edge-on detectors can be employed. One, two or more than two layers of detectors can be employed. Detector modules that comprise a detector layer can optionally be partially-offset or completely-offset from their adjacent neighbors within a layer.

PET image acquisition formats based on planar and focused structure (such as ring and or partial ring) geometries are implemented. Compton-PET detector systems are based on block, 1D, 2D or 3D edge-on, face-on, or mixtures of edge-on and face-on detectors (including edge-on detectors with SAR capability and face-on detectors with DOI capability) (Nelson, U.S. Pat. No. 4,560,882; U.S. Pat. No. 4,937,453; U.S. Pat. No. 5,258,145; U.S. Pat. No. 6,583,420; U.S. Pat. No. 7,291,841; U.S. Pat. No. 7,635,848; U.S. Pat. No. 8,017,906; U.S. Pat. No. 8,115,174; U.S. Pat. No. 8,115,175; and U.S. Pat. No. 8,183,533). The non-coincidence and coincidence Compton-PET detector systems described herein include focused and unfocused planar detector formats and focused structure (such as ring and partial ring as well as focused ring and focused partial ring) detector formats.

A non-coincidence Compton-PET (one-sided PET) detector system is implemented by extending Compton camera designs that have been developed for nuclear medicine imaging devices such as hand held probes or SPECT/Gamma cameras so that the detector system can operate with the lower gamma ray energies used in nuclear medicine as well as the higher energy range of PET with good detection efficiency. A highly flexible implementation of a Compton camera design is a dual-layer, 3D Compton camera. A specific implementation, a non-coincidence Compton-PET detector system, employs a (preferably, but not exclusively) Compton scattering front-end detector and a (preferably, but not exclusively) high-stopping power back-end detector in which both front-end and back-end detectors offer suitable 3D spatial resolution, energy resolution and temporal resolution (Nelson, U.S. Pat. No. 8,017,906).

Both the front-end and back-end 3D detectors provide adequate temporal resolution for an expected event rate, such that accurate event tracking can be enabled both within the front-end and back-end detectors and between the front-end and back-end detectors. Both the front-end and back-end 3D detectors can record Compton scatter and photoelectric interactions.

In some instances Raleigh scattering interactions (angle change with insignificant energy loss) can be identified based on tracking information. The front-end and back-end detectors, either separately or together, can operate as two layer Compton cameras and Compton telescope cameras (Nelson, U.S. Pat. No. 8,017,906).

In one scenario the 3D front-end detector can function as a single (or multiple) Compton scatter device and the 3D back-end detector can be used to measure the energy and interaction location of the Compton scattered photon. The front-end and back-end detectors have 3D spatial resolution. Front-end and back-end 3D detectors can also Compton-scatter a photon (measuring position and energy deposited) and detect the (single or multiple) Compton-scattered photon (measuring its energy and interaction location). Therefore this two layer Compton camera with 3D detector layers incorporates the capabilities of three two layer Compton cameras (in which one layer Compton-scatters the photon and the other layer detects (stops) the Compton-scattered photon).

Compton telescope camera designs exploit multiple Compton scattering for reconstruction. The Compton telescope camera capability can be implemented in the 3D front-end detector, in the 3D back-end detector and between the 3D front-end and back-end detectors (providing the capabilities of three (multilayer, face-on 2D array detectors) Compton telescope cameras).

Appropriate two layer Compton camera and Compton telescope camera reconstruction algorithms are used to form an image. When this Compton camera is used to image single annihilation gamma rays created during a PET scan it is referred to as a one-sided PET detector system or a non-coincidence Compton-PET detector system. (This dual-layer, 3D Compton camera design is clearly not limited to PET imaging alone and therefore may be adapted for use in imaging applications at other photon energies. Furthermore more than two layers of 3D detectors can be employed and non-3D layers of detectors can be mixed with 3D layers of detectors, thereby introducing additional flexibility in the types of imaging applications for which this Compton cameras design is suitable.)

This one-sided PET detector can be implemented in a focused or unfocused planar detector geometry or a focused structure detector geometry such as a ring or partial ring (as well as focused ring and focused partial ring detector geometry). This avoids the expense of employing a coincidence PET detector system based on opposing (or nearly-opposing) sets of PET detectors.

Examples

FIG. 1 shows a dual-layer Compton-PET detector imaging system 1000 that incorporates 3D, edge-on detector arrays 510 and 520 (a first layer of detectors and a second layer of detectors, respectively). The individual, 2D edge-on detector modules 102 use crossed strip radiation detectors 115. Alternatives include 2D pixelated arrays (or 3D pixelated arrays if SAR capability is enabled) in an edge-on geometry.

Incident radiation photons 107 from gamma ray radiation source, with energy less than the pair production threshold, can undergo Rayleigh scattering, Compton scattering or photoelectric interactions. Compton scattered gamma ray photons 108 can be detected by the edge-on radiation detector within the module 102 responsible for the initial scattering, by other edge-on detectors modules within the front-end detector layer 510 (detector layer 1) or by detector modules within the back-end detector layer 520 (detector layer 2).

Each module 102 also includes a base 106 and a communications link 103. The base 106 preferably contains detector electronics including signal conditioners and readout ASICs, power management components, temperature control components, and a data or information channel for communicating with the computer system. The communications link 103 can be used to provide power to the module 102 and connects the base 106 to a computer system.

The communication link 103 preferably is used to off-load the digitized detector radiation data to a computer system for analysis and image reconstruction. The computer system, which can include general purpose, dedicated, and embedded computers, provides monitor and control services to modules 102, to the detector layers 510 and 520 and to the entire Compton-PET detector imaging system 1000.

The computer system evaluates module parameters, detector layer parameters, and the detected radiation image data. The detected data is processed and can be displayed and stored if desired. Additional relevant module information, such as temperature, amplifier settings, detector voltages, position, orientation, and motion information, can be transmitted to this computer system over the communication link 103. The computer system transmits instructions that update the detector modules 102 and detector layers 510 and 520. This establishes a dynamic information feedback loop that is useful for adaptive imaging (Nelson, U.S. Pat. No. 7,291,841).

Note that the electronic functionality of the detector base 106 can be implemented along the side of a detector module or attached to the surface of the detector module (integrated electronics). Another option when the detector substrate is a semiconductor such as Si is to etch an indentation along the bottom of (opposite the radiation entrance surface) and mount the readout ASICs and radiation shielding in the indentation and directly to the substrate along the bottom edge.

If the length of the edge-on detector is greater than its height then this configuration allows the readout ASICs to be closer to a set of detector pixels than for the case wherein the readout ASICs are mounted along the side in order to avoid the direct x-ray beam. Preferably the combined thickness of the etched substrate and mounted readout ASIC with shielding would be less than or equal to the thickness of the substrate (avoiding problems if the readout ASIC and any shielding stick out from the substrate and possibly interfering with the x-ray beam seen by offset detectors).

FIG. 2 shows a perspective view of readout ASICs 200 with radiation shielding 204 mounted in an etched Si substrate 208 (or another suitable semiconductor substrate), with a pixel size 215 that varies with height which is positioned edge-on to incident radiation photons 109. Other means of delivering power to the detector modules as well as wireless communication can be employed in place of communication link 103 (FIG. 1). It should be understood that readout ASICs can be mounted along the side and the bottom edge.

Two or more non-coincidence Compton-PET detector systems (an enhanced non-coincidence Compton-PET detector system) can be employed for a PET imaging application. Furthermore, with the addition of coincidence circuitry, pairs of non-coincidence Compton-PET detector systems (preferably facing each other and positioned on opposite sides of an object) can function as a coincidence Compton-PET detector system.

The cost of a two layer non-coincidence Compton-PET (one-sided PET) detector system can be reduced if either one or both of the 3D front-end and back-end detectors can be replaced by a suitable 2D detector with acceptable energy and temporal resolution and stopping or scattering power. The caveat is that photon detection efficiency and reconstruction image quality may suffer as a result. A compromise in terms of cost is to leave the front-end detector with 3D spatial resolution (and therefore retaining the previously listed capabilities: to function as a Compton scatterer, a two layer Compton camera, a Compton telescope camera), and employ a back-end detector with 2D spatial resolution. The back-end detector would offer acceptable stopping power, energy resolution and temporal resolution for the expected gamma ray event rate and gamma ray energies.

For a planar detector geometry the front-end and back-end detectors can consist of single-layer face-on detector plane modules, a multilayer (stack) of face-on detector plane modules, a single-layer of edge-on detector modules, a stack of edge-on detector modules or a combination of face-on and edge-on detector modules. Face-on detector modules can include DOI capability whereas edge-on detector modules can include SAR capability.

One implementation of a focused planar detector geometry (suitable for cone beam CT, tomosynthesis, etc.) employs a front-end detector that consists of either a single layer (offset or non-offset) or multiple layers (offset or non-offset) of tilted edge-on (and/or face-on) detector modules. A degree of physical focusing (promoting directionality) is achieved by tilting the detector modules (detector modules with fixed or adjustable tilt angles can be implemented, depending on the imaging requirements). As an alternative to a parallel pixel structure a focused pixel structure can be implemented along the lengths of the edge-on tilted (or parallel) detector modules to account for x-ray beam divergence (which can also be implemented in a ring or partial ring CT detector geometry).

Furthermore, an additional degree of physical focusing can be achieved by positioning detector modules (using parallel and/or focused pixel structures) in a curved geometry and thereby approximating arc-shaped detector lines (suitable for a focused, near-planar detector geometry as well as ring or partial ring CT detector geometries). Each of the offset or non-offset edge-on detector module comprising the first layer of tilted edge-on detector modules can have at least a second (offset or non-offset) edge-on or face-on detector module (comprising the at least second layer of detector modules), tilted or not tilted, positioned beneath it. For example, the first layer can implement offset tilted edge-on silicon detector modules with each offset silicon detector module followed by one or more semiconductor or scintillator face-on or edge-on detector modules comprising one or more additional layers (typically employing moderate-to-high Z detector materials).

FIG. 3 shows a perspective view of a focused planar detector system 1000 in which detector modules 102 are tilted so as to focus on diverging radiation 109 from a radiation source. In addition the pixel structure 115 within the individual detector modules 102 is angled so as to focus on the same radiation source.

The tilting of the detector modules may create unacceptable gaps between neighboring detector modules within the detector layer 510. These gaps are shown to be effectively filled by the complete offset of every other detector module comprising the offset detector layer 510.

One implementation of a focused structure detector geometry such as a ring (or partial ring) employs a front-end detector comprised of a single layer (non-offset) or single layer with an offset layer (which can be treated in this application as a single layer) of tilted edge-on detector modules. As in the case of planar detectors, a focused pixel structure can be implemented along the lengths of the edge-on tilted detector modules (creating focused ring and focused partial ring detector geometries).

Suitable detector configurations and materials have been described for Compton, PET, nuclear medicine and x-ray imaging (Nelson, U.S. Pat. No. 6,583,420; U.S. Pat. No. 7,291,841; U.S. Pat. No. 7,635,848; U.S. Pat. No. 8,017,906; U.S. Pat. No. 8,115,174; U.S. Pat. No. 8,115,175; U.S. Pat. No. 8,183,533; U.S. patent application Ser. No. 13/199,612; and U.S. patent application Ser. No. 13/507,659). Examples of suitable detector configurations include a single or multilayer face-on detector, a single or multilayer edge-on detector and a multilayer detector comprised of face-on and edge-on detectors.

Edge-on detectors may incorporate SAR capability and face-on detectors may incorporate DOI capability. Examples of suitable detector materials and formats include semiconductor detectors, structured detectors such as single and double sided structured 3D silicon (Parker S., et al., IEEE Trans. Nucl. Sci. 53, pp. 1676-1688 (2006); Da Via C., et al., Nucl. Instr. Meth A594, p. 7 (2008)) as well as other structured 3D semiconductor materials (Diamond, Ge, Se, GaAs, CdTe, CZT, etc.), structured quantum dots (Urdaneta, M., et al., IEEE Nuclear Science Symposium (oral presentation, 2010)), structured scintillators, and scintillators. Structured mold quantum dot detectors (also referred to as structured quantum dot detectors) offer flexibility since a variety of cell shapes (including trenches) can be implemented (Nelson, U.S. patent application Ser. No. 13/507,659). Furthermore, the selection of (and density of) quantum dot materials can be varied as a function of position within the substrate in order to enhance a type of interaction such as Compton scattering or the photoelectric effect. Silicon is frequently used as a mold material in the form of porous silicon or micromachined silicon for semiconductor quantum dots. Silicon and other mold materials can be used with scintillator quantum dots as well as conventional scintillator materials.

Structured mold semiconductor detectors implement (but are not limited to) either semiconductor quantum dots or amorphous semiconductors or polycrystalline semiconductors (semiconductor materials). The flexibility of the structured mold architecture enables incorporating not only two or more semiconductor materials within a structured mold but also implementations such as one or more semiconductor materials with one or more scintillator materials and/or gases, one or more scintillator materials with one or more gases, etc. within a structured mold. For example, an edge-on dual-layer detector with a semiconductor detector first (front-end) layer and a scintillator detector second (back-end) layer can be manufactured as a single, edge-on structured mold detector with semiconductor and scintillator components.

The first layer within the structured mold could implement one or more semiconductor quantum dot, amorphous semiconductor and/or polycrystalline semiconductor materials in appropriate geometries (in this implementation the first layer is comprised of one or more layers) for the incident radiation field. The second layer could implement one or more organic and inorganic scintillator materials including, but not limited to, scintillator quantum dot, polycrystalline scintillator, nanophosphor scintillator, liquid scintillator, gas scintillator, etc. materials in appropriate geometries (in this implementation the second layer is comprised of one or more layers) for the incident radiation field.

Partial lists of suitable organic and inorganic scintillators and semiconductors are provided, e.g., in Knoll G., Radiation Detection and Measurement, 4th edition, Wiley (2010). Suitable materials include, but are not limited to, organic crystal scintillators, inorganic crystal scintillators, plastic (polymer) scintillators and (plastic and non-plastic) scintillating fibers and fiber bundles (strips) (scintillating fiber bundles (strips) represent one implementation of a structured detector), gel scintillators, liquid scintillators, deuterated liquid scintillators, and loaded liquid scintillators (loaded, e.g., with B, Gd or Sn). Suitable gas scintillators include, but are not limited to, xenon, krypton, argon, helium, and nitrogen. Glass scintillators may also be used (e.g., silicate glass containing lithium activated with cerium).

Additional detector options include structured, gas-filled straw detectors with appropriate low-Z or moderate-Z material annuli which provide suitable energy, spatial and temporal resolution and stopping or scattering power (Nelson, U.S. Pat. No. 8,017,906), liquefied gas based detectors (such as Xenon), semiconductor-based or gas-based Medipix detectors and low temperature (such as GE and superconductor) detectors. Multiple Compton-PET (one-sided PET) views of a volume of an object to be imaged can be acquired as a result of detector system rotation about the object to be imaged.

An alternative imaging format is to rotate the object and keep the detector system stationary. Additional object volumes can be imaged, if needed, by translating (typically) the object through the scanner system.

It should be noted that if the Compton camera image quality isn't suitable for the nuclear medicine imaging applications of interest then a collimator can be inserted in front of the detector so that the system of collimator and detector can function as a SPECT/gamma camera. Since the collimator imposes a degree of directionality then the SPECT/gamma camera implementation of a Compton camera can utilize both Compton scatter interactions (and tracking capabilities) as well as direct photo-electric interactions (which have a much higher probability of occurring at lower energies such as 140.5 keV versus 511 keV in low-Z and high-Z detectors). The direct photo-electric interactions would not be used in conventional (no electron tracking) Compton camera imaging. Furthermore, a miniature version of the Compton-PET detector system can be implemented as a Compton-PET hand-held detector probe. The addition of a nuclear medicine collimator permits the Compton-PET detector probe to function as a gamma camera hand-held detector probe. Versions of probes can be operated in non-coincidence or coincidence mode with non-coincidence Compton-PET detector systems (as well as coincidence Compton-PET detector system) to offer enhanced resolution.

Coincidence Compton-PET detector systems extend the implementations of a non-coincidence Compton-PET detector system by including a second Compton-PET detector system and coincidence circuitry between the pair of Compton-PET detector systems, for example, employing a pair of planar or partial ring Compton-PET detector systems with coincidence circuitry.

FIG. 4 shows a perspective view of a coincidence Compton-PET detector system which is comprised of a pair of planar Compton-PET detector systems 1000 with communications links 103 operated in coincidence for imaging an object 111 (for example, the heart). Each planar Compton-PET detector system 1000 is positioned by an electronically controlled actuator arm 130.

For the case of a partial ring Compton-PET detector system, if a sufficient number of pairs of partial ring Compton-PET detector systems and coincidence circuitry (linking all detectors) are employed, then a complete ring coincidence Compton-PET detector system can be implemented. The complete ring geometry can be achieved with a single pair of partial ring Compton-PET detector systems if each partial ring covers an angular aperture of 180 degrees.

If the Compton scatter capability of a front-end detector is not needed (for example, if only one complete Compton camera is needed for non-PET image applications), then there is the option of employing only a PET-compatible detector for the second detector system. Additional pairs of Compton-PET and/or PET-compatible (or combinations of both) detectors with appropriate coincidence circuitry can be combined to form an enhanced coincidence Compton-PET detector system. (Note that a dummy or non-functional equivalent of the front-end detector can be used to make a stand-alone PET-compatible detector unit “see” a comparable radiation field to what the back-end detector experiences in a coincidence Compton-PET system without the cost of an active front-end detector).

The description of a flexible non-coincidence Compton-PET detector system applies to the Compton-PET detector systems used in a coincidence Compton-PET detector system. Consider the case in which at least one of the two detector system is a Compton-PET detector system. The front-end and back-end detectors offer suitable 3D spatial resolution, energy resolution and temporal resolution and stopping or scattering power. Both the front-end and back-end detectors provide adequate temporal resolution for an expected event rate such that accurate event tracking can be enabled both within the front-end and back-end detectors and between the front-end and back-end detectors, since Compton scatter and photoelectric interactions can be recorded in both front-end and back-end detectors.

As described for non-coincidence Compton-PET detector systems this combination of front-end and back-end detectors incorporates the capability of three two-layer Compton cameras and three Compton telescope cameras. The addition of coincidence detection capability introduces additional flexibility in that events involving a single photoelectric interaction (in which no Compton scattering occurs) in the front-end or back-end detector can be used for coincidence detection as well as events involving one or more Compton scatter interactions.

In a conventional Compton camera design a photoelectric interaction in the front-end detector layer is not useful. A fast or very fast detector (including, but not limited to, silicon, GaAs, structured 3D silicon, structured mold, etc.) can provide timing information for coincidence PET using either photoelectric or Compton scatter interactions. The capability to use photoelectric events without Compton scattering leads to an alternative detector system design in which the front-end detector layer employs a moderate-to-high Z detector material with tracking capability. In this implementation both photoelectric events and Compton scattering can be exploited but now the photoelectric interaction relative probability is more significant compared to a material such as silicon. Tracking capability for Compton scattered photons (as well as characteristic x-rays) can be used for estimating the deposited energy for each detected event even if Compton scatter reconstruction is not employed. If the front-end detector is sufficiently fast then the TOF PET imaging can be implemented. For example, an edge-on, structured mold detector implementing at least one of high-Z (semiconductor) quantum dots, amorphous semiconductors, and polycrystalline semiconductors can replace an edge-on silicon or structured 3D silicon detector. If the edge-on, structured mold detector offers significant attenuation it can be used in place of both the front-end and back-end detectors. Furthermore, this alternative detector system design can be readily extended for use with CT-Compton-PET, CT-PET, PET Compton-PET and CT detector systems.

Since very fast coincidence timing (TOF) can be used to improve reconstruction accuracy and reduce patient dose and/or image acquisition time there can be a benefit from having one or both of the front-end and back-end detectors capable of very fast timing resolution. If both front-end and back-end detectors are involved in the detection process then coincidence timing can be based on using at least one of the front-end and back-end interaction timings. Timing resolution corrections are made for the response of one or both detectors (depending on whether one or both of the front-end and back-end detectors are involved in detection) and gamma ray travel times between interaction locations within one or both detectors and between detectors (Nelson, U.S. Pat. No. 8,017,906).

Commercial TOF PET systems are capable of very fast temporal resolution (on the order of or less than one nanosecond). Very fast temporal response capabilities can influence the choice of detector materials for front-end and back-end detectors. If the front-end detector has a reasonable probability per photon of a Compton scatter interaction then one option is to select a front-end detector material with a very fast temporal response and select a (possibly much less expensive) back-end detector material with a much slower temporal response.

If a gamma ray undergoes a Compton scatter interaction in at least one of the front-end and back-end detectors as well as additional interactions such that the energy of the incident particle can be estimated, then photon directionality based on the appropriate Compton camera reconstruction algorithm (for the Compton camera designs described for non-coincidence Compton-PET detector system) can be compared with photon directionality based on coincidence (line-of-sight) between the Compton-PET detector systems operating in coincidence. The Compton-based directionality can be used to estimate the degree of validity of the coincidence (line-of-sight) assumption, including acollinearity. This capability can be used to help reject some of the photons that undergo Raleigh and/or Compton scattering within the object and its surroundings as well as Rayleigh scattering or difficult to detect Compton scattering within the detectors.

In addition, a (combined) non-coincidence Compton-PET (one-sided PET) reconstructed image can be compared to (or combined with) a coincidence PET reconstruction image. (Nelson, U.S. Pat. No. 8,017,906). Unpaired detected events (in which coincidence fails since only one of the two annihilation photons is detected and is considered legitimate) by a Compton camera can still contribute to the Compton scatter reconstruction image.

As described for the case of non-coincidence Compton-PET (one-sided PET) detector systems, system cost (in some cases) may be reduced if the back-end detector 3D spatial resolution capability is lowered to 2D capability while maintaining adequate energy and temporal resolution. The 2D spatial resolution of the back-end detector implies that it offers limited performance as a stand-alone PET detector for gamma rays that aren't Compton scattered by the front-end detector.

The back-end detector should provide good stopping power. The Compton scattering front-end detector offers suitable 3D spatial, temporal and energy resolution and scattering interaction probability. Single and multiple Compton scattering (as well as photoelectric) interactions can occur in the front-end detector, allowing the front-end detector to function as a Compton camera, as a PET camera, as the first layer of a multilayer Compton camera and as the first layer in a multilayer PET camera in which it records the initial interaction location, energy deposition and event timing information. (Note that if the multilayer Compton camera capability is sacrificed then the 2D spatial resolution capability of the back-end detector can be reduced to 1D or even block detector spatial resolution, further reducing costs. The back-end detector primarily provides stopping power along with appropriate energy and temporal resolution. The front-end detector should offer an acceptable probability of undergoing at least one Compton scatter interaction so that an initial location of interaction, timing and energy deposition can be established. If TOF PET imaging is desired then the front-end detector can offer very fast temporal resolution. The front-end detector, due to its 3D spatial resolution capability, can still track multiple scatter interactions as well as photoelectric events. The front-end detector retains the capabilities of a Compton camera and a PET detector. Event tracking between the front-end and back-end detectors is employed.)

The back-end detector can also offer fast or very fast temporal resolution. The front-end detector can maintain fast or very fast temporal resolution capability but an alternative is to implement a front-end detector (with suitable temporal resolution) primarily for establishing spatial resolution via photoelectric and Compton interactions while relying on the fast or very fast back-end detector (with suitable spatial resolution) to establish coincidence timing for the front-end detector Compton scattered photons. In some implementations cost savings may be realized and the choice of front-end and back-end detectors may be expanded by moving some temporal resolution capabilities from the front-end detector to the back-end detector. For example, the 3D front-end detector could use low, moderate or high Z semiconductor materials with less emphasis on temporal resolution and more emphasis on spatial resolution while the back-end detector could offer reduced spatial resolution while emphasizing temporal resolution (employing fast scintillators such as LSO, LYSO, BaFl₂, LaBr₂, etc. coupled to PMTs, microchannel plates, SiPMs, etc.). Implementations of these dual-layer detector configuration are suitable for use in CT-PET imaging systems wherein the front-end detector is used for CT and PET and the back-end detector is used for CT and PET or PET. Furthermore, the front-end detector and back-end detector layers can always be implemented as independent detector layers leading to cost savings. For example, the CT front-end detector can be a simple face-on scintillator (energy integrator), a dual-energy edge-on scintillator, a PCE edge-on semiconductor detector, etc. but the needed for capabilities such as tracking electronics between layers, etc. are removed.

Multiple Compton-PET or PET views of an object volume to be imaged can be acquired as a result of detector rotation about the object. The alternative imaging format is to rotate the object and keep the detector system stationary. If the Compton camera image quality isn't suitable for the nuclear medicine imaging applications of interest then a collimator can be inserted in front of the detector so that the system of collimator and detector can function as a SPECT/gamma camera (collimators can also be used with PET cameras).

For the coincident and non-coincident Compton-PET configurations described there are many options for detector materials based on cost and performance requirements. Assuming that acceptable-to-good energy resolution is desirable, then block, 1D, 2D and 3D back-end detectors and 2D and 3D front-end detectors can use semiconductors, polycrystalline and amorphous semiconductors, structured 3D semiconductors, structured mold semiconductor quantum dots (nanoparticles) as well as amorphous semiconductors and polycrystalline semiconductors, moderate-to-bright nanophosphors (scintillator quantum dots), organic and inorganic scintillators, gas and liquid detectors, and amplified detectors. Furthermore these detectors can incorporate edge-on SAR or face-on DOI (positional encoding) capabilities.

Semiconductor and gas detectors typically offer a Fano factor noticeably less than 1.0. If stopping power is important, then sufficient detector material can be present in order to provide good to excellent attenuation. Detector response time (for example, scintillator decay time) properties should be suitable for at least event tracking at expected event rates. Very fast detectors would permit the use of TOF information to be utilized in PET reconstruction algorithms.

Possible scintillators with at least one of these properties include, but are not limited to: BaFl₂, LaBr₂, LaCl₂, LSO, LYSO, GSO, GdI₃, LuI₃, SrI₂, BaHfO₃, SrHfO₃, PbWO₄, LuAP, CsI:Tl,Sm, NaI:Tl, BGO, CsI:Tl, Lu₂O₃:Eu, ZnO-based fast scintillators as well as glass, plastic and fiber scintillators, liquid scintillators, gas scintillators, quantum dot scintillators, ceramic scintillators, polycrystalline scintillators and various fast-to-very fast organic scintillators. Possible semiconductor detectors with at least one of these properties include, but are not limited to: diamond, Si, SiC, Se, Ge, GaAs, CdTe, CZT, HgI₂, PbO, PbI₂, TlBr (as well as low noise implementations such as silicon drift detectors or those with gain such as Si-APDs or SiPMs or iDADs, Se-APDs, GaAsPMs and DiamondPMs) detectors; one dimensional structures such as rods and wires, structured single and double sided 3D Si and other semiconductor material detectors (Parker S., et al., IEEE Trans. Nucl. Sci., 58, pp. 404-417 (2011)) and structured mold semiconductor quantum dot, amorphous semiconductor, and polycrystalline semiconductor detectors.

A number of these semiconductor detectors can be configured as fast or very fast photodetectors and so they can be coupled with fast or very fast scintillators such as quantum dot, organic, or inorganic scintillators. Suitable detector packages (a detector material coupled to a readout ASIC) include Medipix-based detectors. Additional structured detectors with gain include, but are not limited to, gas-filled straw detectors (Nelson, U.S. Pat. No. 8,017,906).

In addition, the choice of detector material can be influenced by the detector format. For example, a 10 mm thick (or greater) CdTe or CZT face-on detector (used primarily for photo-detection) for PET imaging may offer limited temporal resolution, whereas a 1 mm thick (or less) CdTe or CZT edge-on detector (used for photo-detection and/or Compton scattering) may qualify as a fast detector (even if SAR or DOI corrections are not implemented). From a similar perspective a 1 mm or 0.5 mm (or less) thick Si or Ge edge-on detector (used for Compton scattering or Compton scattering and photo-detection) can be employed as a fast or very fast detector.

If SAR or DOI capabilities are implemented to estimate the interaction location of an event, then timing corrections can be made based on the propagation times of electrons or holes to the anode and cathode, respectively (Nelson, U.S. Pat. No. 7,635,848; U.S. Pat. No. 8,017,906). An edge-on or face-on structured 3D semiconductor or structured mold semiconductor quantum dot detector can be employed as a fast or very fast detector since charge propagation distances can often be less than 40-100 microns.

The flexibility of this Compton-PET design also allows alternative choices for the front-end detector and back-end detector based on factors such as lower cost and non-redundancy of features (if possible) as well as spatial resolution, energy resolution, temporal resolution and the likelihood of Compton scatter and photoelectric interactions. For example, a Compton-scatter front-end scintillator detector could be employed based on suitable (excellent) timing resolution despite reduced energy resolution compared to a semiconductor detector. Suitable front-end detector candidates with at least one of these properties include low-Z or moderate-Z, fast and very fast organic or inorganic scintillators (or scintillating fibers) with a suitable high-speed, sensitive optical readout detectors (such as photodiodes, APDs, semiconductor photomultipliers such as SiPMs and GaAsPMs, electron multiplier CCDs, microchannel plates, etc.), semiconductor-based, scintillator-based or gas-based Medipix detectors, and structured, gas-filled straw detectors with appropriate low-Z or moderate-Z material annuli (including the straw material itself) which function as a source of Compton electrons. In additional examples, straw and plastic (or non-plastic) scintillating fiber detector formats, when implemented in an edge-on geometry used for PET imaging, could implement SAR (permitting position estimates as well as energy and timing corrections). Furthermore, plastic scintillating fibers (as well as non-plastic scintillating fibers) can be coated with thin films of moderate-to-high-Z materials to enhance their photoelectric cross section (permitting the properties of the front-end detector layer to be tuned in terms of the Compton scatter and photoelectric interaction probabilities).

Previously, structured straw detectors incorporated only high-Z annuli in order to enhance the photoelectric effect (Nelson, U.S. Pat. No. 8,017,906). The same design technique can be used with low-Z and moderate-Z annuli in order to enhance the Compton scatter effect. Furthermore, combinations of low/moderate-Z annuli straw detectors followed by high-Z annuli straw detectors (or other high-Z detectors) can be employed.

Detectors should offer an acceptable probability of experiencing at least one Compton scatter interaction so that an initial location of interaction can be established. Event tracking within and between the front-end and back-end detectors can be employed. If the front-end detector offers excellent temporal resolution then TOF information can be used to improve the reconstructed image along with a reduction in patient dose and/or image acquisition time. If a front-end detector lacks good energy resolution it still can be effective if the front-end and back-end detectors offer good spatial resolution and the back-end detector offers good energy resolution.

Coincidence (line-of-sight or line-of-reaction) directionality can be exploited along with the scattered photon angle in order to estimate the incident gamma ray energy for cases of simple Compton scatter. Once the properties of the front-end or back-end detector have been defined, then the properties of the other detector can be selected on the basis of which properties need to be accentuated or can be allowed to diminish (such as stopping power, energy resolution, spatial resolution and temporal resolution).

The back-end detector may primarily offer stopping power and energy resolution if the front-end detector offers 3D spatial resolution and energy resolution. Then a cost-based decision can be made as to whether the front-end or back-end detector (or both) should provide acceptable, fast or very fast temporal resolution.

Thus a single detector implementation does not have to embody all of the coveted PET detector properties (high stopping power, 3D spatial resolution, fast or very fast temporal resolution). For example, the coincidence Compton-PET detector system can implement features such as TOF imaging with a range of detector options that is much greater than with commercial (conventional) TOF PET systems. Nonexclusive lists of suitable scintillator and semiconductor materials are provided herein. Partial lists of suitable organic and inorganic scintillators and semiconductor materials including some of their properties are provided in Knoll G., Radiation Detection and Measurement, 4th edition, Wiley (2010), p. 230 (table 8.2), p. 238 (Table 8.3) and p. 492 (Table 13.3), respectively, which are incorporated by reference herein.

The flexibility of using front-end and back-end detectors for PET which can offer different spatial, temporal and energy resolution for PET results in different PET images based on which detectors interact with the pair of gamma rays from an annihilation event. For example, a Compton-PET front-end detector could Compton scatter one gamma of a pair which is then detected by the back-end detector. Another Compton-PET front-end detector might fail to scatter the other gamma of the pair which is detected by the back-end detector. Coincidence can be established but the timing or spatial resolution (or both) of the front-end detector that detects one gamma may be much better than the timing or spatial resolution of back-end detector that detects the other gamma of the pair.

The use of front-end and back-end detectors permits flexibility as to which detector parameters to optimize (temporal, spatial, energy resolution) as well as detector material properties (density, Compton scatter versus photoelectric interaction probability, Compton or photoelectric electron range), for the front-end and back-end detectors. Cost-sensitive decisions can made based on detector characteristics and geometries in terms of how they influence various PET parameters including energy resolution, spatial resolution, temporal resolution, sensitivity, NECR (noise equivalent count rate), true counts, incorrectly classified events, random events, characteristic radiation, Rayleigh scatter, acollinearity, etc.

For example, it may be suitable to employ 0.5 mm thick, high-resistivity or detector grade pixelated silicon or a structured (3D) silicon (or structured mold) detector arranged edge-on (for adequate energy resolution, improved spatial resolution, faster timing), rather than 1.0 mm thick, detector grade Silicon arranged face-on. Or a material with a higher Z than Silicon could be employed to increase photoelectric interaction probability (Ge, GaAs CdTe, CZT, structured 3D and structured mold detectors, etc.). One possibility is that a front-end detector alone will be adequate. For a dual-layer (or multilayer) detector system all detector interaction combinations (and thus a range of PET images with different properties) need to be considered.

Consider a dual-layer detector in which the two layers may have one or more different properties such as stopping power, spatial resolution and timing resolution. The first layer could, for example, be comprised of an array of edge-on, high spatial resolution (typically small pixels), fast or very fast temporal resolution, (low-Z) silicon or structured silicon 3D detector planes (or structured mold detectors) providing 3D detector capability. The second layer could be an array of edge-on or face-on, moderate or high-Z, semiconductor or scintillator or structured 1D or 2D detector planes (providing 2D or 3D detector capability, respectively) of the same or lower spatial and temporal resolution (typically slower, larger pixels).

Photoelectric interactions that occur in the first layer or second layer as well as valid reconstructed events (the result of tracking of single scatter or multiple scattered photons as well as shared energy within or between layers) can be used in coincidence detection with an opposing dual-layer detector. Note that this will result in multiple PET images. Coincidence between opposing (typically faster, smaller pixels) first layer detectors (based on photoelectric events or tracked scatter events interacting within the first layer) may be best for spatial and timing resolution information followed by coincidence between a first layer detector and an opposing second layer detector. The poorest spatial and timing resolution would be provided by coincidence between (typically slower, larger pixel) second layer detectors.

As in the case of a dual-layer or “telescope” Compton camera which employs tracking the synergistic interaction of the detector layers enables the recovery of a fraction of the scatter events that interact within one or both layers. A low-Z semiconductor material such as silicon, in which Compton scattering dominates the photoelectric effect at 511 keV, can be used in high resolution PET imaging since the photoelectric and scattering effects can be exploited. If Compton reconstruction algorithms can be employed the effects of acollinearity and scatter may be reduced in some instances and some non-coincidence detected events can be used to form a non-coincidence PET image (Compton-PET image). If advantageous, data from one or more types of acquired PET images can be combined to reconstruct enhanced PET images. Note that low-Z semiconductor detectors such as silicon as well as structured mold semiconductor detectors and other structured semiconductor detectors (for example, structured 3D silicon detectors) are also suitable for use in PET detector systems as well as CT detector systems (ring, partial ring, cone beam) in single layer or offset layer detector formats.

CT-Compton-Pet Detector Systems

The flexibility of the Compton camera design allows it to be adapted for PET (and nuclear medicine) imaging. The Compton camera design can also be adapted for use in diagnostic x-ray imaging applications such as CT and projection radiography (with the understanding that typical data rate requirements will be much higher, spatial resolution requirements may increase, and the operational energy range for diagnostic medical CT is typically lower than for PET and nuclear medicine imaging).

Various coincidence and non-coincidence Compton-PET detector system implementations have been described. An extension of this dual-use concept is to describe a multi-use CT-Compton-PET detector system design (with the understanding that nuclear medicine imaging capability can also be implemented).

The incorporation of CT features can be explained by examining a special case of a Compton-PET detector system design, the CT-Compton-PET detector system design. This is of interest because CT-PET detector imaging systems are commercially available. However, the CT and PET detector imaging sub-systems (which use face-on detectors) are physically distinct. This commercial configuration involves moving the patient with respect to the typical partial ring geometry (or alternatively a cone beam geometry) CT scanner into a physically separate PET scanner. These CT and PET detector sub-systems do not share detectors or the image acquisition space.

An alternative to the existing commercial CT-PET detector imaging systems are improved CT-PET detector systems in which the CT scanner or PET scanner (or both) are replaced with novel edge-on CT scanners and/or PET scanners (including Compton-PET detectors) described in this application. For example, the face-on detector CT configuration is replaced with an edge-on CT detector system capable of performing at least one of energy integration, PC, and PCE (Nelson, U.S. Pat. No. 6,583,420; U.S. Pat. No. 7,291,841; U.S. Pat. No. 7,635,848; U.S. Pat. No. 8,017,906; U.S. Pat. No. 8,115,174; U.S. Pat. No. 8,115,175; and U.S. Pat. No. 8,183,533).

For example, one CT implementation would employ a single layer using an array of edge-on semiconductor detectors operating in PCE mode (such as an edge-on semiconductor silicon detector or an edge-on structured semiconductor detector). Suitable edge-on structured semiconductor detectors include at least one of a 3D semiconductor detector such as 3D silicon, a structured mold semiconductor detector incorporating one or more of semiconductor quantum dots, amorphous semiconductors, and polycrystalline semiconductors. If additional stopping power is needed a second layer of moderate-to-high Z semiconductor or scintillator detectors could be implemented.

One implementation of a second detector layer is to employ face-on or edge-on semiconductor arrays (including structured detectors) which operate in PCE mode (based on factors such as cost, detector response limitations, and/or information content of the radiation field). An alternative (less-costly) implementation is for the second layer to employ a face-on or edge-on scintillator array operating in integration mode or PC mode to primarily detect the more energetic x-rays (providing additional information about the radiation field to compliment the spectral information acquired with the first detector layer). If advantageous, detectors can implement mode switching circuitry (for example, from PCE mode to integration mode or PC mode) as a means of compensating for excessive event rates.

A fast, improved CT-PET detector system incorporates multiple x-ray tubes (two, three or more) or x-ray sources (such as carbon nanotubes, scanning electron beams, etc.) to reduce image acquisition times. Novel PET detectors include, but are not limited to, 3D crossed rod, crossed fiber-rod and encoded PET detectors. The physically separate PET or Compton-PET scanner preferably provides one or more detector features such as suitable or excellent energy resolution, 3D spatial resolution and TOF capability. If reduced PET performance is acceptable then one or more of energy, spatial and temporal resolution can be degraded.

PET designs described in this patent application can be employed with commercial face-on CT scanners to comprise enhanced CT-PET detector systems. Physically separate commercial PET scanners can also be used with an edge-on CT detector system in another version of an enhanced dual CT-PET imaging system. Still another version of an enhanced dual CT-PET imaging system employs physically separate edge-on CT and PET designs described in this application and prior patents. Yet another version of an enhanced CT-PET imaging system is to employ a face-on detector or edge-on detector CT scanner with a physically separate Compton-PET detector system.

Alternative to commercial and enhanced dual CT-PET detector designs are CT-Compton-PET systems in which detector components and/or space are shared, representing a cost effective and compact design compared with the benefit that the patient remains stationary and so registration between CT and PET images is straightforward. Furthermore current CT imaging sub-systems in commercial dual CT-PET systems do not offer PC or PCE capabilities which are available in enhanced dual CT-PET and CT-Compton-PET detector systems.

PC or PCE capabilities can be used for dose reduction and/or multispectral imaging. Furthermore, multispectral imaging can be implemented with a PC detector system by implementing x-ray tube voltage switch (currently employed with dual-energy CT detector systems).

CT-Compton-PET detector systems designs incorporate the capabilities described for Compton-PET detector systems. One or more layers of detectors can be employed. PET options include non-coincidence (one-sided) and coincidence PET imaging capabilities. The incorporation of x-ray CT capabilities may impose additional requirements on the design of the radiation detectors, depending on the energy range for the application (small animal, pediatric, adult, therapy, industrial, HLS, synchrotrons) and the event (data) rates (which, for medical CT imaging, are typically much higher than the event rates encountered in nuclear medicine imaging).

In addition, collimation may be introduced into the CT detector which would be of a relatively fine nature. The type and amount of collimation introduced into the CT detector configuration is preferably sufficient to at least result in a beneficial reduction in radiation cross talk between detector elements during CT imaging without substantially reducing the efficiency of the PET detector component of the imaging system. If external collimation is employed to reduce the intensity of x-rays scattered by the object from reaching the CT detector, and this external collimation has an undesirable impact on PET imaging efficiency or image quality, then the external collimation is preferably moveable so that it can rotate or slide out of the detector field of view (FOV) during PET imaging.

X-ray scatter correction algorithms in CT imaging can also be employed with or without collimation along with corrections for detector effects such as induced charges in nearest-neighbor detector elements, charge cloud diffusion and radiation cross talk (energetic electrons, characteristic x-rays, bremsstrahlung) between detector elements (Nelson, U.S. Pat. No. 7,291,841; U.S. Pat. No. 8,017,906). If the PET detector imaging is not implemented simultaneously with the CT detector imaging then an optional movable, attenuating shield (such as, but not limited to Cu, W, Pb, a multilayer material) can be inserted during CT imaging to protect the PET detector from unnecessary radiation damage, and then removed during PET imaging.

The insertion of nuclear medicine collimation hardware such as parallel or pin hole (pinhole) collimators into these Compton camera designs can provide nuclear medicine imaging capabilities for those cases in which the Compton camera does not offer adequate imaging properties. CT detector modes of operation can include energy integration, PC or PCE. One implementation of a CT-Compton-PET detector system is to simply operate the back-end PET detector independently of the front-end CT detector and accept that the CT detector acts as an attenuator and scatterer of the 511 keV PET gamma rays.

More sophisticated CT-Compton-PET detector systems will be described next. Implementations of detector geometries include planar (and focused planar) configurations and focused structure configurations such as rings and partial rings (as well as focused rings and focused partial rings). Planar, ring, and partial ring detector geometries are encountered in medical diagnostic x-ray CT.

CT-Compton-PET detector system designs described herein are based on implementations of coincidence and non-coincidence Compton-PET detector systems with additional constraints imposed by CT imaging. X-ray fluence rates for diagnostic medical x-ray CT are typically sufficiently high that features such as PC and PCE are easier to implement if the distribution of detected events during a time interval is spread out over a number of detector channels. Other constraints on detector selection are related to problems such as dose-dependent pixel performance degradation (including polarization issues) and detector effects described herein.

This tends to limit the selection of edge-on or face-on detector to one or more fast-to-very fast, low-to-moderate Z semiconductors with or without gain capability (including, but not limited to, Si, Ge, GaAs, diamond, Se, Si-APDs, SiPMs, iDADs, Se-APDs, GaAsPMs, DiamondPMs), structured 3D semiconductor detectors and structured mold (quantum dot, amorphous, polycrystalline) semiconductor detectors coupled to high speed readout circuitry (such as a custom readout ASIC or a Medipix chip). Other options include configurations such as gas-based Medipix detectors and fast-to-very fast, bright scintillators coupled to photodetectors.

Other semiconductor material such as CdTe or CZT may be employed if they are sufficiently thin (typically less than 1 mm) such that issues related to polarization at high data rates can be mitigated. Their pixel performance degradation rates and detector effects should be acceptable (or can, in part, be compensated by evaluating whether any correlated charge was deposited in neighboring pixels as in the case of the Medipix detector chip).

For the case of a focused structure detector geometry such as a ring the detector modules can form partial rings, with detectors in a single partial ring that have small gaps or gaps comparable in thickness to 2D edge-on detector plane modules (with optional collimation between the detector plane modules). If gaps are of comparable thickness to the 2D edge-on detector plane modules then the x-ray source is preferably collimated to match the gaps in the detector plane and the collimators and detector need to move along the ring by one pixel width (detector plane width) to acquire a complete projection for reconstruction. This compensating motion and matching x-ray source collimation is not needed if at least two sets of partially-offset or completely-offset detector rings (alternate detector modules are located at two different radii) with gaps comparable to the thickness of 2D edge-on detector modules are employed (Nelson, U.S. Pat. No. 7,291,841).

The CT edge-on detector modules employed in a focused structure ring geometry can also be employed in a planar CT detector geometry. One or more layers of edge-on detector modules can be configured to be parallel or tilted with respect to adjacent detector modules in order to achieve a focusing effect. As with the ring geometry implementations, layers of tilted edge-on detector modules can also be partially offset or completely offset so that tilted edge-on detector modules in a lower layer fill gaps between edge-on detector modules in the upper layer(s) so that a reasonably continuous detector is emulated.

As described, a focused pixel structure can be implemented along the lengths of the edge-on tilted (or parallel) detector modules. Various configurations of edge-on or face-on (single or multilayer) detector modules or combinations of face-on and edge-on detector modules may also be employed in planar and ring detector geometries. Optionally, SAR and DOI capabilities can be incorporated into the edge-on and face-on detector modules, respectively.

If the front-end CT detector and the back-end Compton-PET (or PET) detector operate independently of each other, then the CT-Compton-PET detector system can be considered a limited CT-Compton-PET detector system (an integrated limited CT-Compton-PET detector system). In this case the range of front-end CT detector designs extends from planar to focused structure (ring and partial ring) geometries and from traditional (low-cost) energy integration detectors to PC to PCE detectors.

The front-end CT detector attenuates a fraction of annihilation gamma rays directed toward the back-end Compton-PET (or PET) detector. The planar or focused structure back-end Compton-PET (or PET) detector does not have to occupy the same FOV as the CT detector; larger or smaller FOVs can be implemented according to hardware constraints, cost and desired acquisition times. The back-end Compton-PET (or PET) detector can be designed to operate with 2D or 3D spatial resolution.

Non-coincidence PET (one-sided PET) imaging can be implemented with a limited CT-Compton-PET system in which the back-end detector is a Compton-PET detector. For coincidence PET imaging, the back-end Compton-PET (or PET) detector can provide either 2D or 3D spatial resolution capability.

Coincidence PET imaging may require the addition of a second PET detector system and the appropriate coincidence circuitry. If the Compton-PET detector offers 3D resolution and tracking capability then both coincidence and non-coincidence PET imaging can be conducted simultaneously. Another implementation of a limited CT-Compton-PET detector system is to position the Compton-PET (or PET) detector outside the FOV of the CT detector. A radiation shield may be inserted between the CT detector and the Compton-PET (or PET) detector during CT operation to limit unnecessary radiation dose to the Compton-PET (or PET) detector system.

For example, consider the case where at least one CT detector array (also referred to as a CT detector) is employed (high speed acquisition may require multiple distinct CT detector arrays and x-ray sources or even a full ring CT detector array). An implementation for a ring-like acquisition geometry can employ at least one front-end (inner layer) CT detector as a partial ring detector (or a planar detector) for PET (preferably) aligned with a pair of back-end (outer layer) opposing partial ring (or planar) PET detectors or a full ring PET detector. Optionally, a fraction of the CT detector can be implemented with PET detection features.

For the case of a dedicated PET or Compton-PET imaging system the one or more partial ring (or planar) CT detector arrays or a full ring CT detector array can be replaced with comparable (or smaller) detector arrays with the advantage that the pixel geometry and readout electronics can be optimized for PET or Compton-PET imaging without consideration for CT pixel geometry or CT readout electronics. Note that in an alternative implementation at least one front-end CT detector array with a planar format can be aligned with a pair of back-end opposing planar or partial ring PET detectors or a full ring (or square/rectangular) PET detector. One or more (aligned or unaligned) additional front-end (AFE) detectors (not necessarily used for CT) can be incorporated into the detector geometry in order to improve PET imaging system capabilities such as detection efficiency and/or timing resolution.

For example, an AFE detector could be paired with the CT detector (or a fraction thereof). Extra AFE detectors can be added, either positioned independently or positioned as opposing pairs. The case of a single front-end partial ring CT detector aligned with a pair of back-end opposing partial ring PET detectors permits many interactions to be considered (the partial ring CT detector can also interact with partially-aligned or unaligned PET detectors). A Compton image can be acquired by using the CT detector for scattering gammas, a fraction of which are then preferably detected by either the aligned PET detector behind it or by the CT detector alone (e.g., if it has 3D capability). Another option is for the PET detector alone (if it has 3D capability) to scatter and detect incident gammas that did not interact with the CT detector. The information from Compton scattering a sufficient number of times within or between detectors can be used for Compton image reconstruction based on multiple scattering.

Alternatively, the information from Compton scattering one or more times terminating with a photoelectric event can be used for Compton image reconstruction. Compton image reconstruction can also be implemented in the opposing, aligned PET detector (e.g., if it has 3D capability). If a Compton image event can also be used for PET coincidence imaging then the Compton image information may be used to enhance the PET image since additional information concerning the directionality of the detected gamma is available. PET coincidence images can be acquired based on coincidence between a PET detector with a PET detector, a CT detector or a CT/PET detector (a CT/PET detector represents, e.g., a CT detector interacting with preferably an aligned back-end PET detector although the CT detector can also interact with partially-aligned and unaligned back-end PET detectors).

If at least one AFE detector is present then new coincidence images can be acquired including AFE detector coincidence with a CT detector (and AFE/PET if the AFE can interact with a PET detector), a PET detector, a CT/PET detector and other AFE (as well as AFE/PET) detectors (if present). Furthermore, AFE (and AFE/PET, if present) Compton images can be acquired. Image data can be combined when appropriate to synthesize enhanced diagnostic images. The dual-layer detector formats described herein can also be implemented as dedicated PET detectors (or nuclear medicine/PET detectors) including the benefits of generating multiple types of coincidence images and/or Compton images, using Compton image information to enhance coincidence PET imaging and the synthesis of enhanced diagnostic images. Note that each layer in the dual-layer detector formats described can be comprised of sub-layers (and/or have a detector with a structure that is equivalent to having sub-layers). Furthermore, detector properties within a layer or sub-layer can vary (a design feature also applicable for dedicated multilayer and single layer PET systems).

For example, consider the front-end partial ring CT detector used with a back-end PET ring detector with detectors that offer only moderate temporal resolution (not necessarily suitable for TOF PET). The front-end partial ring CT detector typically offers high spatial resolution but it may or may not offer high temporal (fast or very fast) resolution. If high temporal resolution for the PET image associated with the CT detector is desired and the CT detector is not fast then the aligned PET detector segments behind and opposite the CT detector should preferably implement suitably fast detectors (e.g., replacing both of the moderate temporal resolution PET detector segments with fast PET detector segments). Alternatively, if the CT detector is fast then the aligned PET detector segment opposite the CT detector should be suitably fast whereas the aligned PET detector behind the CT detector can be fast but it is not required to be fast (if the interaction probability of 511 keV gammas with the front-end CT detector is satisfactory for generating sufficient coincidence events with the opposing PET detector segment during image acquisition). Then one or both moderate temporal resolution PET detector segments may be replaced with suitably fast or very fast detector segments.

Since the back-end PET detector segments are expected to detect a fraction of unscattered gammas another option is to implement fast PET detector segments regardless of the temporal response of the front-end detectors. Yet another option is to introduce at least one AFE detector that may or may not offer high temporal resolution to operate with the opposing PET detector segment permitting increased flexibility with respect to properties implemented for the opposing PET detector segment. The PET detector segments for CT-Compton-PET or dedicated PET systems can optionally be designed to include positioning capability permitting greater flexibility for image acquisition optimization. Cost can factor into the decision if the fast PET detector segments cost substantially more than the PET detector segments they would replace. Other tradeoffs such as differences in detector stopping power and readout electronics requirements need to be considered.

For CT-Compton-PET detector systems the front-end CT detector also serves as front-end detector layer for a Compton-PET detector system. The readout electronics should be suitable to handle event data rates that are on a comparable scale to the event data rates experienced by CT detectors, or the CT detector pixel geometry can be modified to reduce the effective data rate per pixel and so reduce the requirements of the readout electronics. The front-end and back-end detector layers preferably include appropriate internal and intra-layer event tracking capabilities (for coincidence and non-coincidence Compton-PET systems) depending on their intended use.

For CT applications which utilize PC or PCE capabilities several edge-on pixel geometries have been described including uniform pixel sizes (1D or 2D pixel array) and non-uniform pixel sizes (Nelson, U.S. Pat. No. 7,635,848; U.S. Pat. No. 8,115,174; U.S. Pat. No. 8,115,175; and U.S. Pat. No. 8,183,533). Issues arise as to x-ray beam hardening with depth of penetration and the benefit of imposing a more-uniform distribution of interaction rates between pixels along the x-ray beam direction (reducing readout errors and readout electronics costs).

If the event rate is sufficiently low a uniform pixel distribution may be adequate even if beam hardening occurs with penetration depth. If the event rate is high (as expected in many diagnostic medical x-ray CT applications) and PC or PCE capability is required, then a static, uniform 2D pixel array may not offer a good balance in detected event rate per pixel unless the pixel dimensions are relatively small in terms of the stopping power of the detector material. Implementing such relatively small pixels allows a degree of flexibility since a variable effective pixel size versus depth could be synthesized by combining the output signals from two or more pixels.

Unfortunately, as pixel size decreases the number of pixels and readout electronics increases which raises the cost of the detector modules. In addition to detector effects pixel readout noise can increase due to leakage issues associated with some small pixel implementations.

High event rates and x-ray beam hardening with penetration depth may favor the use of a non-uniform pixel size with increasing detector depth along a pixel column. The pixel length within a column can increased with increasing depth, resulting in a non-uniform (variable) readout element pitch in order to provide a more-balanced count rate per pixel for the readout electronics.

Detector pixel distributions as well as the use of collimating septa and/or side shielding for detector modules used in CT systems have been described (Nelson, U.S. Pat. No. 6,583,420; U.S. Pat. No. 7,291,841; U.S. Pat. No. 7,635,848; U.S. Pat. No. 8,017,906; U.S. Pat. No. 8,115,174; U.S. Pat. No. 8,115,175; and U.S. Pat. No. 8,183,533). Furthermore, the pixel size in the axial direction (the slice thickness) can be non-uniform (benefitting dose reduction). For example, a high resolution pixel size (thin slices) could be implemented near the center of the detector in the axial direction with a lower resolution pixel size (thicker slices) implemented on both sides of the center.

Additional non-uniform pixel size distributions can be implemented based on imaging requirements. Additional flexibility is provided when the outputs of two or more pixels in the axial direction can be combined electronically in order to synthesize the desired distribution of pixel sizes in the axial direction. A non-uniform pixel size in the axial direction can be implemented with edge-on detectors and face-on detectors. A non-uniform pixel size distribution can be implemented along an arc segment.

The high spatial resolution detectors can (in one implementation) be positioned at the middle of the detector arc (that images the region of interest within the object being scanned) with low spatial resolution detectors on either side. With edge-on detectors the low spatial resolution detectors can be synthesized by combining the outputs of two or more pixels with the same coordinates as measured with respect to the edge-on detectors themselves. Thus, comparable pixels from adjacent edge-on detectors (even if they are offset with respect to their neighbor) are combined.

Both PCE and PC readout modes can be deployed as needed according to the imaging requirements along the axial direction and along the arc (such as the need for energy subtraction in a limited region of image). Alternatively, an integration readout mode can be implemented if a PCE or PC readout mode is or will be saturated. Appropriate beam collimation and filtration can be employed to match the pixel distribution in the axial direction and along the arc. Furthermore, non-uniformity can be extended to include the detector geometry type (mixing of edge-on and face-on detectors). For example, high spatial resolution edge-on detectors are (in one implementation) positioned at the middle of the detector arc (that images the region of interest within the object being scanned), with low spatial resolution face-on detectors on either side (potentially reducing the over-all cost of the detector system).

The principles of non-uniformity in pixel size and detector geometry type can be applied to both ring and planar detector systems. Detector configurations of reduced size can be employed if region of interest CT is implemented (retaining the high spatial resolution detectors that image the region of interest within the object being scanned while eliminating the low spatial resolution detectors on either side).

A focused structure, ring geometry Compton camera design (Nelson, U.S. Pat. No. 7,291,841), may or may not offer optimal performance as a CT-Compton-PET camera for high event (data) rate, fan beam CT diagnostic imaging. The Compton camera would preferably use edge-on detector modules with a uniform pixel size along a column (uniform 3D spatial resolution), whereas the PC or PCE CT system would preferably use edge-on detector modules with a variable readout element pitch along a column.

The variable readout element pitch for CT allows the readout rate requirements of the readout ASIC-based electronics to be better balanced between readout elements (pixels) near the entrance surface and pixels distant from the entrance surface, which experience reduced beam intensity. Thus the number of readout elements can be reduced noticeably and fewer readout ASICs of a given performance level are needed compared to a uniform pixel array with many small pixels. If the readout ASICs electronics offer high readout data rates sufficient to handle the maximum expected CT data rates for any pixel in a uniform pixel detector which is preferred for use in a Compton camera or Compton-PET detector, then this not an insurmountable constraint.

A potential drawback is a likely increase in cost due to a need for more high speed readout ASICs than would be utilized in a dedicate CT scanner with similar PC or PCE capabilities, but a non-uniform pixel distribution with depth. Other issues that may arise due to this CT-COMPTON-PET detector system design and the increased use of high speed readout ASICs are related to an increase in heat generation and therefore new cooling requirements to avoid increased detector noise and thermal expansion issues. There is also a possibility that some readout ASICs may be moved closer to the pixels (which may result in certain readout ASICs positioned within the x-ray beam path and therefore altering shielding requirements).

Note that this issue of CT detectors with uniform and non-uniform pixel arrays in CT-Compton-PET detector systems affects both the focused structure ring (or partial ring) detector format used in fan beam CT and the planar detector format used in cone beam CT. One alternative is to use readout ASICs of varying performance with respect to depth. The highest speed readout ASICs would read out the pixels close to the entrance surface, whereas readout ASICs of progressively slower speeds (but still sufficient for both CT and Compton camera applications) could be used to read out pixels at greater depths.

Another alternative is to enable the edge-on detector module electronics to redefine the readout element pitch according to whether the CT-Compton-PET detector system is functioning as a PET detector system or a CT detector system. Thus, a detector module can have a selectable (fixed or variable) effective pixel width along a detector row and/or an effective pixel length along a detector column in which the effective pixel width or length is synthesized from the outputs of one or more (typically) uniformly spaced pixels.

For example, a variable, effective pixel length can be optimized for CT imaging based on the beam spectrum and the beam intensity. A softer x-ray beam would preferentially be attenuated closer to the detector entrance surface than a harder x-ray beam, for a given detector material (for energies away from a detector material k-edge). For the case of a softer x-ray beam of a given intensity the balancing of event rates between successive effective pixels in a column would benefit from electronically synthesizing relatively smaller effective pixel lengths near the entrance surface. Relatively larger effective pixel lengths would create a better balance of event rates between effective pixels in the case of a harder x-ray beam of a given intensity.

The advantage of a synthesized readout is that it can be optimized according to the energy spectrum and the desired readout rates, thus expanding the use of a PC or PCE CT system to a broad range of beam spectrums (applications) while retaining the uniform detector pixel geometry useful for PET (and Compton camera) imaging. Since a SPECT camera employs collimation to define directionality of the incident photons, either uniform or non-uniform detector pixel geometry can be employed (making a CT-SPECT detector system relatively straightforward to implement with appropriate collimation in place).

If tracking of Compton-scattered photons within the SPECT camera is implemented, then a uniform detector pixel geometry may be beneficial. Features such as redefining the readout element pitch (synthesizing an effective pixel length or width) or employing readout ASICs of varying performance with detector depth can be implemented in dedicated CT detector systems, as well as CT-Compton-PET detector systems and CT-SPECT detector systems. Furthermore, CT-SPECT detector systems can employ a single detector layer or multiple detector layers.

CT-Compton-PET detector system geometries include planar and focused planar detector systems and focused structure detector systems such as ring and partial ring (as well as focused ring and focused partial ring) detector systems. Non-coincidence and coincidence CT-Compton-PET configurations are described herein based on non-coincidence and coincidence Compton-PET configurations. The CT x-ray detectors offer suitable 3D spatial resolution, energy resolution (PCE capability) and temporal resolution to be useful for the high x-ray fluence rates encountered in medical and non-medical CT scanning as well as for use as the front-end detector in non-coincidence and coincidence Compton-PET detector systems. Event tracking capability may be required for CT-Compton-PET systems.

Non-coincidence CT-Compton-PET detector systems combine CT imaging capability with one-sided PET imaging capability by employing the CT x-ray detector as the front-end detector layer that would be used in a non-coincidence Compton-PET detector system in conjunction with a high-stopping power back-end detector. A flexible design employs front-end and back-end detectors that offer suitable 3D spatial resolution, energy resolution and temporal resolution.

Both the front-end and back-end detector layers provide adequate temporal resolution for an expected event rate, such that accurate event tracking can be enabled both within the front-end and back-end detectors and between the front-end and back-end detectors, since Compton scatter and photoelectric interactions can be recorded in both front-end and back-end detectors. All implementations described for non-coincidence Compton-PET (three two-layer Compton cameras and three Compton telescope cameras) are applicable, possibly with the added constraint that the front-end detector should offer suitable detection efficiency for the x-ray energy spectrums that would be used in CT imaging, should be compatible with the event rates for CT imaging, and should offer a spatial resolution with depth that is reasonably uniform when Compton and/or PET imaging modalities are employed.

FIG. 5 shows a perspective view of a CT-Compton-PET detector system 1000 in a focused structure (partial ring) geometry which includes a point-like x-ray 109 radiation source 125 and a gamma ray 107 radiation source 111. The front-end detector layer 510, comprised of detector modules 102 which use 2D pixelated array radiation detectors 115 in an edge-on geometry with base 106 and communication links 103, performs the dual role as an x-ray CT detector and a front-end detector layer (detector layer 1) for a Compton-PET detector system.

The detector modules 102 are mounted in a rigid structure 110. The back-end detector layer 520 (detector layer 2) could be of a planar or focused structure geometry. For comparison, FIG. 1 can be understood to show the front-end and back-end detector layers 510 and 520 (detector layers 1 and 2) for a planar CT-Compton-Pet detector geometry if the front-end detector layer 510 is suitable for CT imaging.

A reduction in cost can be realized if the Compton-PET capability is implemented only within a sub-region of the CT detector (for example, a segment of a partial ring detector geometry or a region of a planar detector geometry). In these instances segments of CT detector modules or regions of CT detector modules that are not involved in PET imaging do not need to implement features such as synthesizing variable effective pixel lengths or employing readout ASICs of varying performance with detector depth. Multiple Compton-PET views can still be acquired as a result of detector rotation (in some applications the object can rotate and the detector is stationary).

By reducing the active detector area the detection efficiency will be reduced and acquisition times will, in general, increase. Alternatively, acquisition times can be typically be reduced by increasing the PET detector FOV beyond the CT detector FOV. As described, if the Compton camera image quality isn't suitable for the nuclear medicine imaging applications of interest then a collimator can be inserted in front of the detector so that the system of collimator and detector can function as a SPECT/gamma camera.

Coincidence CT-Compton-PET detector systems extend the implementations of non-coincidence CT-Compton-PET detector systems with the addition of coincidence detection capability by introducing a second Compton-PET detector system along with appropriate coincidence circuitry. If the Compton scatter capability of a front-end detector is not needed then only a PET-compatible detector may be needed for the second detector system.

Implementations described for coincidence Compton-PET detector systems are applicable. Thus, the detector geometries shown in FIG. 1 and FIG. 5 are applicable when employed in a coincidence detection configuration such as FIG. 4. Again, a reduction in cost can be realized if the coincidence Compton-PET or coincidence PET capability is implemented only within a sub-region of the CT detector (for example, a segment of a partial ring or complete ring detector geometry or a region of a planar detector geometry) and a matching Compton-PET or a PET-compatible back-end detector of comparable dimensions is positioned opposite that segment or region of the CT detector.

Additional cost savings may be realized if the second coincidence Compton-PET detector system employs a front-end detector that offers comparable performance to the CT detector when used as part of a Compton-PET detector system, but lacks the extreme performance capability of a CT detector. Acquisition times can be typically be reduced by increasing the PET detector FOV beyond the CT detector FOV.

Multiple Compton-PET or PET views of a limited volume of the subject can be acquired as a result of detector rotation about the subject. In some applications the subject can rotate and the detector is stationary. By reducing the active detector area detection efficiency may be reduced and acquisition times may increase. If the Compton camera image quality isn't suitable for the nuclear medicine imaging applications of interest, then a collimator can be inserted so that the detector can function as a SPECT/gamma camera.

The CT-COMPTON-PET scanner assigns the CT detector to the role of a front-end detector in a Compton-PET detector system when Compton camera or PET (or nuclear medicine) imaging is implemented. In an implementation of a coincidence Compton-PET detector, the front-end detector primarily acted as a Compton scatterer (with photoelectric detection capability) and the back-end detector provided stopping power, energy resolution and temporal resolution sufficient for event tracking with respect to the front-end detector. Options described for the front-end detector include, but are not limited to, sufficiently thin planar semiconductor detectors, structured 3D semiconductor detectors, structured mold quantum dot semiconductor detectors, detectors with SAR or DOI capability, low/moderate-Z scintillator detectors and structured low/moderate-Z straw detectors (which typically require lower data rates than the semiconductor-based detectors).

Furthermore, the front-end CT detector may be a multilayer detector, as described herein. For example, one implementation employs a front-end detector comprised of a first layer of an edge-on semiconductor followed by an edge-on or face-on scintillator second layer followed by a back-end PET detector which now functions as a third layer (a variation of this detector design employs a suitably-designed back-end PET detector to function in the role of the second layer of the front-end CT detector). (Similar detector implementations may be used for the back-end detector although the detector properties may differ between front-end and back-end detectors.)

If the front-end detector offers an acceptable Compton interaction probability with annihilation gammas and it is fast enough to provide the required coincidence timing resolution (or very fast coincidence timing if TOF PET imaging is desired), then the back-end PET detector requirements can be simplified since its role is primarily to detect (typically through the photoelectric effect) Compton-scattered gammas from the front-end detector. If the back-end detector is required to provide coincidence resolution (including TOF resolution if desired), then the selection of suitable detector materials and detector designs may be reduced. In one implementation the back-end PET detector is also used to detect annihilation gammas that don't interact with the front-end detector (requiring coincidence resolution or TOF resolution). (Front-end and back-end detectors can function independently as PET coincidence detectors, front-end and back-end detectors can function cooperatively as a PET coincidence detector, and front-end and back-end detectors can function cooperatively as a Compton gamma camera.)

Reduced spatial resolution could be acceptable for a back-end detector (although Compton camera reconstruction accuracy will be reduced or lost) used in coincidence PET imaging, if the front-end detector provides adequate 3D Compton-scatter information including moderate-to-high spatial resolution. For example, a single detector block, a 1D or a 2D detector array could be implemented based on factors such as expected count rate, required energy resolution, cost and desired flexibility.

In general, for both coincidence and non-coincidence Compton-PET detector systems, a combination of a 3D back-end detector with a 3D front-end detector could improve overall detection efficiency. In this implementation the back-end detector could detect Compton scattered photons from the front-end detector, unscattered (primary) photons using the photoelectric effect, and photons scattered within the back-end detector itself. Face-on and edge-on (or angled) detector designs can be employed for the back-end detector as well as the front-end detector.

PET scan times can be improved by employing additional partial-ring or planar PET or Compton-PET detector systems that operate with or are independent of the coincidence or non-coincidence CT-Compton-PET detector system. These systems are referred to as enhanced coincidence or non-coincidence CT-Compton-PET detector systems. The amount of rotation about the object to acquire a more-complete PET image can be reduced.

Another option is to implement a coincidence CT-Compton-PET detector system based on a multiple (two or more) x-ray tube or x-ray source system. For example, the angular arc of a commercial, dual x-ray tube CT partial ring detector is approximately twice that of a single x-ray tube system. Multiple cone beam imaging can be implemented if there are two or more x-ray tubes or x-ray sources and corresponding planar detectors. (An example of a multi-planar detector/x-ray tube CT system developed for high speed cardiac and lung CT was the Mayo Dynamic Spatial Reconstructor or DSR first implemented in the late 1970s.) Note that if interior tomography techniques can be implemented, then x-ray intensities and/or areas of planar detectors (depending on the application) may be reduced (Yu, H. and Wang, G., Phys. Med. Biol., Vol. 54(9): pp. 2791-2805 (2009)).

For the case of the focused structure partial ring geometry the CT partial ring detector (the front-end detector) used in a dual x-ray tube configuration can be split into two equal CT partial ring detector sections so that at least one CT partial ring detector section (and its back-end detector) can be rotated through 180 degrees when coincidence PET scanning is initiated. This could be particularly beneficial for applications such as fast scan Cardiac CT in conjunction with Cardiac PET CT. Other applications that could benefit from high resolution CT and PET or SPECT (nuclear medicine) imaging capabilities of this system include head imaging and small animal imaging. Note that the back-end detector might cover only a segment of a CT partial ring or complete ring detector (or a region for a planar detector). If coincidence CT-Compton-PET system is implemented the second planar or partial ring PET detector may only need to be comparable in size to the actual PET detector implemented with the first CT planar or partial ring detector.

The efficiency of a PET detector system can be improved by adding additional front-end detectors (and corresponding back-end detectors) opposite to, adjacent to or separate from the CT partial ring detector or the CT planar detector. These front-end detectors could utilize less demanding readout electronics and may not require features such as pixel synthesis since they would only be used for PET imaging and not CT imaging. Note that for the various PET implementations in which an opposing PET detector would block the x-ray beam path the opposing PET detector is either rotated out of the beam path (the x-ray tube or x-ray source may be physically retracted when not in use) or a small opening is made in the opposing PET detector to pass the collimated x-ray beam (the PET detector rotates with the x-ray tube or x-ray source).

Multiple x-ray tubes or x-ray sources (as described for fast, improved CT-PET detector imaging systems) can be employed with enhanced integrated non-coincidence or enhanced coincidence CT-Compton-PET detector imaging systems and enhanced limited integrated CT-Compton-PET detector imaging systems. Both stationary and rotating x-ray tube-detector systems can be implemented (both designs have been used with dedicated CT imaging systems).

Dedicated (stand-alone) CT detector imaging system in a ring or planar detector geometry can be implemented by reducing the functionality of the CT-Compton-PET detector imaging systems described herein. As detailed, detectors with fixed (or variable) uniform or non-uniform pixels can be implemented with the requirement that the detectors can perform efficiently at the event count rates per pixel encountered in medical CT imaging.

CT detectors include single layer and multilayer detectors comprised of face-on detectors and/or edge-on detectors including gas, scintillator, semiconductor, low temperature (such as Ge and superconductor) and structured detectors (such as structured 3D semiconductor, structured mold quantum dot and scintillator-photodetector structured detectors). Single layer and multilayer detector designs of Compton cameras described herein can be implemented in a dedicated CT detector imaging system with PCE capability (a simplification would be a design that provides PC capability). Multilayer designs typically maintain or increase the atomic number of the detector material for progressively deeper detector layers with respect to the radiation entrance surface.

Consider a single layer, edge-on detector implementation for a medical CT imaging system in which detector planes are aligned with the Z-axis in a ring geometry. 2D Si edge-on detectors with a wafer thickness of (for example) approximately 500 microns (μm) as currently implemented may be preferred over relatively thick, expensive, face-on CdTe or CZT detectors in terms of operational lifetime and temporal response. Alternative edge-on detectors of comparable thickness (approximately 500 microns) which can offer improvements with respect to the stopping power and/or temporal response performance of 2D Si at reduced cost compared to the relatively thick, face-on CdTe or CZT detectors include, but are not limited to, 2D ZnO (which is also a fast, relatively low-Z scintillator making ZnO attractive as a semiconductor or scintillator detector material for TOF PET), 2D GaAs, 2D CdTe and 2D CZT detectors (as well as low noise implementations, implementations with gain or low temperature implementations such as 2D Ge) and structured detectors (structured 3D semiconductor detectors such as 3D Si, 3D GaAs, 3D CdTe, 3D CZT, 3D Ge, etc., as well as structured mold (quantum dot, amorphous, polycrystalline) semiconductor quantum dot detectors).

If cost is an issue and reduced capabilities (such as reduced energy resolution) are acceptable then a structured mold scintillator quantum dot detector could be employed (for example, functioning as an energy integrator detector for single or multiple energy CT). Since quantum dot density can be varied with position and a single quantum dot material or multiple quantum dot materials can be employed as a function of position it is readily apparent that the equivalent of a multilayer detector can be synthesized within a single structured mold quantum dot detector by varying quantum dot density and/or material as a function of position in regular or irregular patterns, as described below.

This concept of varying material (and/or material density) is readily extended to structured mold amorphous and polycrystalline semiconductor detectors, structured mold scintillator detectors, etc. This design can be further generalized to structured mold semiconductor detectors that include two or more of quantum dots, amorphous semiconductors, and polycrystalline semiconductors, as well as structured mold detectors that incorporate (for example) semiconductors and scintillators, semiconductors with gases, scintillators with gases, etc.

In addition, detectors with thickness greater than or less than 500 microns can be implemented depending on the image resolution requirements for the CT detector imaging system (medical diagnostic, dental panoramic imaging, radiation therapy, industrial, Homeland security, etc.). This single layer, edge-on detector CT imaging system can be employed as a single layer PET imaging system and/or a Compton camera/nuclear medicine imaging system.

As described, multiple Compton-PET implementations are possible. Furthermore, PET and Compton camera/nuclear medicine imaging can be conducted simultaneously. Depending on the fraction of the ring circumference covered by edge-on detectors, additional detectors (of the same or different design) may need to be added to increase coincidence detection efficiency.

For the relatively small (hardware) pixel sizes employed in current medical CT imaging systems, Si is a reasonably efficient detector for the lower x-ray energies encountered in mammography CT and pediatric CT. For adult CT the efficiency of Si may suffer, particularly for x-ray energies above (approximately) 40 keV. A compromise, multilayer detector configuration (for example) could employ edge-on, 2D semiconductor or structured semiconductor detectors (such as 3D Si or GaAs detectors or structured mold semiconductor detectors with materials such as semiconductor quantum dots, amorphous semiconductors, polycrystalline semiconductors) as the low-Z or moderate-Z front-end detector, with moderate-Z or high-Z, edge-on or face-on, back-end detector. (Note that if temperature requirements can be met then Ge is a candidate as a moderate-Z, face-on or edge-on detector.) The back-end detector, with appropriate capabilities, may not only improve the overall CT detector performance but also may be suitable for a different imaging modality such as PET (as described herein). A simplification is to implement a single layer or offset detector layer format for CT or PET using at least one of edge-on, 2D semiconductor detectors, structured semiconductor detectors (such as 3D semiconductor detectors or structured mold semiconductor detectors), and structured scintillator detectors.

Consider the case of a multilayer (in this case, a dual-layer) detector configuration in which an edge-on, 2D Si front-end detector (alternatives such as 3D silicon, etc. may also be employed) is employed as the first detector layer. It would be of reduced height compared to a single layer, edge-on, 2D Si detector implementation and thus less expensive as well as reducing the pixel count and limiting the maximum pixel size. The back-end, second detector layer (edge-on or face-on) is typically comprised of a moderate-Z material (including semiconductors such as GaAs or CdTe or CZT, scintillators and/or structured detectors), or a high-Z material (including semiconductors, scintillators, and/or structured detectors) which would emphasize photoelectric interactions with the high energy photons that penetrate the front-end detector.

One or more types of back-end, face-on detectors can be configured as 1D detectors that are positioned beneath each of the 2D Si edge-on detectors. The thicknesses of appropriate face-on detectors should not be so great that detrimental effects such as polarization or light losses (for scintillators) cannot be mitigated. The cost of manufacturing such 1D detectors (material yields, butting pixels, bonding to readout electronics) should be reduced relative to 2D detectors. More than one layer of 1D, face-on detectors can be employed and layers can consist of the same or different materials. Furthermore, if enhanced detector performance is desired (one or more of: higher spatial resolution, higher temporal resolution, higher energy resolution) in the back-end detector, a 2D face-on detector can be implemented even if SAR is not implemented in the front-end edge-on detector (as noted previously, the detector layers need not have the same spatial, temporal or energy resolution). If 3D information is desired then DOI capability can be introduced or additional layers can be added. Note that in some applications (e.g., due to issues that may be related to cost, temporal resolution, spatial resolution, energy resolution) it may be desirable to employ face-on 1D or 2D (or 3D) detectors in the front end and edge-on detectors (1D, 2D or 3D) in the back end. The useful information that can be extracted from the radiation detected within each layer (as well as cost) will determine whether individual detectors operate as energy integrators, PCs or PCEs.

An alternative is to position a back-end, edge-on 1D or 2D detector (including structured 3D semiconductor and structured mold semiconductor quantum dot, amorphous semiconductor and polycrystalline semiconductor detector implementations) below each front-end, 2D Si edge-on detector. The edge-on, 1D detector is less-costly to manufacture whereas the 2D array will typically handle higher data rates and offer better energy resolution. This dual-layer CT design could be used for both low energy and high energy imaging applications. Any combination of suitable edge-on detectors including 2D detectors, structured 3D semiconductor detectors and structured mold semiconductor detectors such as semiconductor quantum dot detectors can be employed for the front-end and back-end detectors. In this case the semiconductor quantum dots function as semiconductor detector materials and therefore a structured mold semiconductor quantum dot detector can also be described as a specific implementation of a structured mold semiconductor detector.

It should be noted (as described previously herein) that a single layer implementation based only on an edge-on structured 3D semiconductor detector or a structured mold semiconductor detector of a single low, moderate or high-Z material may be implemented in place of a dual-layer CT design. Structured mold semiconductor quantum dot detectors offer additional flexibility, beyond simply varying the dimension of the pixel versus depth to control count rates, in that the density of quantum dots can be varied from low-to-high for individual pixels that constitute the structured quantum dot detector. Another technique to vary quantum dot density is to vary the number of holes within a pixel that are filled with a quantum dot material. For example, pixels could be of a uniform dimensions while energy-dependent attenuation could increase with depth by increasing the density of quantum dots with depth. Thus, the quantum dot (detector) material can be varied as a function of depth in an edge-on orientation. Similar functionality can be implemented with structured mold amorphous and polycrystalline semiconductor detectors. As described herein, structured mold detectors can incorporate more than one type of semiconductor material as well as mixtures of detector material types (semiconductor, scintillator, gas, superconductor).

Furthermore, the selection of back-end detector materials is not limited to semiconductors or structured detectors. The back-end detectors can be face-on or (1D or 2D) edge-on scintillator detectors (Nelson, U.S. Pat. No. 7,291,841). (Additional face-on detector layers can be implemented as needed and different scintillator materials can be employed in different layers.) In general, the back-end detectors can operate as PC, PCE or integrating detectors depending on the application. For example, a PC or integrating back-end scintillator detector can be paired with a PCE front-end detector simplifying detector design. The filtered beam spectrum reaching the back-end detector can be estimated.

In addition, the flexible design permits either or both front-end and back-end detectors to be scintillator-photodetector detectors. For example, the (first layer) front-end detector could be a low-to-moderate Z scintillator-photodetector detector with a (second layer) back-end moderate-to-high Z structured mold quantum dot detector. As noted, the detectors in each layer can operate as either PC or PCE or energy integrating detectors. Thus, one implementation would use a low-Z Si detector with PCE capability for the first layer with a high-Z scintillator-photodetector detector with energy integration capability for the second layer. Furthermore, the flexible design described for dual-layer detector systems can be implemented for multilayer detector system with three or more layers. The single layer and multilayer detector systems described herein can incorporate one or more non-detector materials including attenuating materials, scattering materials and conversion materials depending on the interacting radiation field (e.g., particle types, energies).

It should be noted that a single layer implementation based only on an edge-on structured 3D semiconductor detector or a structured mold semiconductor detector (implementing at least one of semiconductor quantum dots, amorphous semiconductors, and polycrystalline semiconductors) of a single low, moderate or high-Z material may be implemented in place of a dual-layer CT design. Structured mold semiconductor detectors offer additional flexibility, beyond simply varying the dimension of the pixel versus depth to control count rates, in that the density of semiconductor materials such as quantum dots can be varied from low-to-high for individual pixels that constitute the structured mold semiconductor detector. For example, pixels could be of a uniform dimensions while energy-dependent attenuation could increase with depth by increasing the density of semiconductor quantum dots with depth.

Structured Mold Detectors

Structured mold semiconductor quantum dot detectors may deploy a single semiconductor quantum dot material. The use of edge-on, structured mold semiconductor quantum dot detectors creates an opportunity to implement a more flexible detector design. For example, multiple semiconductor quantum dot materials can also be deployed such that low-Z/moderate-Z semiconductor quantum dot materials are positioned near the radiation entrance surface and moderate-Z/high-Z semiconductor quantum dot materials are positioned further from the radiation entrance surface (a multilayer structured mold quantum dot detector). Thus, the selection of semiconductor quantum dot materials can be selected for different energy ranges and the count rate per pixel as a function of distance from the radiation entrance surface can be more-balanced. Furthermore, as noted, the densities of each of the multiple semiconductor quantum dot materials can be varied for individual pixels from low-to-high for optimization purposes.

The semiconductor quantum dot materials can be distributed in appropriate patterns for the incident radiation field utilized for imaging and the modified radiation field within the detector (examples include Compton cameras, spectral CT, etc.). For example, a geometric pattern such as a series of partial concentric rings can be employed to create a focused edge-on detector (with the ability to vary semiconductor quantum dot material and density within a ring and between rings). Furthermore, the partial concentric rings can be comprised of offset pixels (gaps between neighboring pixels in a partial concentric ring that are covered by offset pixels in a neighboring partial concentric ring) rather than a continuum of pixels (FIG. 3 demonstrates a similar design where the gaps between offset pixels in the upper edge-on semiconductor detector layer are covered by the offset pixels in the lower edge-on semiconductor detector layer).

Other semiconductor detector materials can also be employed to fill a 1D array of channels or a 2D array of holes of a conductive mold material, including, but not limited to, polycrystalline and amorphous semiconductor detector materials (e.g., silicon is frequently used as a mold material, in the form of porous silicon or micromachined silicon). When appropriate, both channels and holes can be present in the conductive mold material. In an edge-on orientation the detector material density distribution can be adjusted as a function of depth if desired. For example, a non-uniform detector material density distribution can be implemented to compensate for beam hardening with depth for an edge-on CT detector.

Adjusting the density distribution of a semiconductor detector material within the detector can be implemented in conjunction with pixel size adjustments to provide even greater flexibility in detector optimization for an incident radiation field. Thus, semiconductor detector materials (including semiconductor quantum dot materials) can be distributed according to appropriate patterns for the properties of the incident radiation field (types of particles, energies, angular distribution, intensity distribution, etc.) utilized for imaging. A specific implementation (but not the only possible implementation) of SAR involves segmenting the holes or channels into at least two parts such that separate signals can be read out from the segments.

Structured mold detectors that employ a single detector material (such as a single quantum dot semiconductor material, a single quantum dot scintillator material, etc.) are basic structured mold detectors. Hybrid structured mold detectors can be comprised of multiple detector materials of a single type (for example, multiple semiconductor materials or multiple scintillator materials) or multiple types of detector materials (for example, one or more semiconductor materials combined with one or more scintillator materials and/or gas materials, etc.). An example is a hybrid structured mold semiconductor detector that is comprised of a mixture of semiconductor quantum dot materials and/or other semiconductor materials such as polycrystalline and/or amorphous semiconductor detector materials where low-to-moderate Z materials are positioned to intercept the x-ray beam in the front end while moderate-to-high Z materials are positioned in the back end of the edge-on detector. Furthermore, hybrid structured mold detectors can also incorporate structured 3D detectors such that a region of the structured mold detector is nonporous (the structured 3D detector region) and another region is porous (with pores incorporating detector materials). For example, low-Z 3D silicon is positioned in the nonporous front end of the edge-on detector and moderate-to-high Z semiconductor materials are positioned in the porous back end of the edge-on detector.

Furthermore, basic structured mold detectors and hybrid structured mold detectors can incorporate one or more non-detector materials including attenuating materials, scattering materials and conversion materials to provide transverse filtering and/or lateral shielding. Thus, basic structured mold semiconductor detectors and hybrid structured mold semiconductor detectors can incorporate attenuating materials, scattering materials and conversion materials. For example, patterns of holes or channels can be filled with materials that contain (e.g., for the case of x-ray and gamma ray radiation) moderate-to-high Z attenuating elements (iron, copper, lead, tungsten, uranium, gold, etc.) or alloys in order to create collimation for the radiation field incident on the detector, to reduce inter-detector cross talk elements (lateral shielding). Scattering and conversion materials can be employed in order to alter the properties of the incident radiation field prior to being detected (transverse filtering). For example, conversion materials can be used to transform x-rays to electrons, neutrons to photons, fast neutrons to thermal neutrons, etc. for the benefit of the detector. Scattering materials could be used, for example, to degrade photons or neutrons with undesirable energies. The densities of attenuating, scattering and converting materials can also be varied with position in order to optimize detector performance.

Although the advantages of basic structured mold (and hybrid structured mold) semiconductor detectors have been described for an edge-on detector orientation it should be apparent that similar advantages can be realized for basic structured mold detectors and hybrid structured mold detectors employed in a face-on orientation (or tilted orientation). For example, semiconductor detector materials can be distributed in appropriate patterns for the incident radiation field utilized for imaging, and semiconductor detector material densities can be varied as needed within a single face-on detector layer or with multiple face-on detector layers. Furthermore, patterns of holes or channels can be filled with attenuating materials, etc. in a face-on basic structured detector or a face-on hybrid structured detector.

For example, one implementation would create an attenuating grid pattern providing lateral (side) shielding between face-on detector elements. Non-detector materials such as attenuators, scatterers and converters can also be applied in appropriate patterns on the surface of the face-on basic structured detector or hybrid structured detector (this represents an alternative to introducing a structured layer that only implements appropriate patterns of non-detector materials prior to a structured detector layer). Note that an edge-on SAR capability becomes a face-on DOI capability.

The benefits of basic structured detectors and hybrid structured detectors are not limited to the implementation of semiconductor detector materials, these benefits are also available for implementations where scintillator detector materials (or other detector materials such as gas and low temperature detectors) are employed. Thus, scintillator materials can be arranged in patterns, scintillator density can be varied, multiple scintillator materials can be employed, scintillator and non-scintillator detectors can be employed in the same edge-on or face-on basic or hybrid structured detector and non-detector materials can be employed.

The positioning of detector materials and density distribution of detector materials (as well as non-detector materials) within edge-on and face-on basic structured detectors and hybrid structured detectors can be optimized according to the properties of the radiation field being detected (including mixed radiation fields). In a layered detector system either edge-on detectors or face-on detectors or both edge-on and face-on detectors (as well as tilted detectors) can be employed. Furthermore, both structured detectors and conventional detectors can be employed within a layered detector system. For example, in a two layer system an edge-on, basic structured semiconductor quantum dot detector could be aligned with and followed by a conventional 1D scintillator array coupled to a photodetector (or amplified photodetector) array.

Structured mold semiconductor and scintillator quantum dot detectors (as well as structured 3D detectors and 2D semiconductor detectors) can be implemented with fixed or adjustable pixel sizes which can be uniform or non-uniform. Furthermore, the density of quantum dot material can be varied with position. Typically the lowest density of quantum dot material could be positioned near the radiation entrance surface.

A moderate-Z or high-Z structured mold semiconductor quantum dot detector can also be employed in a face-on orientation as a 1D (or 2D) detector positioned after a (for example) low-Z, 2D Si edge-on detector. Furthermore, moderate-Z or high-Z, fast, bright scintillator-photodetector 1D (or 2D) array detectors (including structured scintillator and nanophosphor detectors), face-on or edge-on, can be employed after a (for example) low-Z, 2D Si edge-on detector (providing limited energy resolution or simply providing photon counting or integration capability or acting as an energy integrators).

The photodetector is a fast, sensitive 1D photodetector that can be chosen from (but is not limited to) photodiodes, APDs, SiPMs, GaAsPMs, DiamondPMs, electron multiplier CCDs and microchannel plates with a pixel structure or a dual-readout structure. Scintillator-photodetector detectors can employ scintillator screens, deposited scintillator films, ceramic scintillators and cut scintillator sheets. Scintillator-photodetector structured detectors can employ structured scintillators (such as manufactured scintillator arrays, scintillators that demonstrate columnar growth and scintillators coupled to fiber arrays) as well as scintillating or minifying scintillating, focused or unfocused, fiber arrays.

Scintillating fiber materials include, but are not limited to, phosphors, granular phosphors, nanophosphors and scintillator quantum dots. If limited energy resolution is acceptable or only photon counting is needed for CT then a moderate-Z or high-Z, fast, bright, scintillator-photodetector or scintillator-photodetector structured detector (where the photodetector is a fast, sensitive photodetector) can be used in place of the single layer or dual-layer detector implementations as described herein (see also Nelson et al., U.S. Pat. No. 4,560,882; U.S. Pat. No. 5,258,145; U.S. Pat. No. 8,017,906; and U.S. patent application Ser. No. 13/507,659).

FIG. 6 shows a minifying scintillating fiber array 140 coupled to a 1D photodetector 141 which is incorporated into the base unit 106. The scintillating fiber array coupled to a photodetector readout comprises a structured detector that can be deployed in place of an edge-on detector in a CT scanner. Adjacent structured detectors such as this can be positioned in a continuous, partially-offset or completely offset configuration.

This ring detector geometry comprised of an array of 1D scintillator-photodetector detectors oriented parallel to the axial direction can be extended to multiple pixel widths along the circumference, since planar or shaped entrance surface scintillating fiber optic arrays and small, 2D high speed photodetector arrays are available. The use of 1D (or 2D) scintillator-photodetector detectors may offer advantages since manufacturing costs are typically reduced, although butting of 1D detectors is generally easier than butting of 2D detectors.

The same approach applies to a planar geometry concerning the use of 1D or 2D scintillator-photodetector detectors. Although readout electronics such as ASICs can be attached to the 1D or 2D photodetector sensors externally, the readout electronics can alternatively be integrated directly on the substrate of the 1D or 2D photodetector sensors.

Orientation, Interaction Height and Sub-Aperture Resolution

Consider a scenario in which radiation is incident upon a planar edge-on detector. The detector thickness (height) now defines the maximum detector entrance aperture while the length or width of the detector area now defines the maximum attenuation distance for edge-on radiation detector designs including semiconductor drift chamber, single-sided strip, and double sided strip detectors, including micro-strip detector versions.

The interaction position along the height of the edge-on detector aperture will be referred to as the interaction height. When a scintillator, semiconductor, gas, or liquid detector is irradiated face-on the 1D positional information along the thickness direction of the detector is referred to as the interaction depth. The electronically-measured face-on detector DOI positional information defines the edge-on detector sub-aperture resolution (SAR).

Strip widths can be tapered or curved if focusing is desired. In the case of double-sided parallel strip detectors in which opposing strips are parallel, both electrons and holes can be collected to provide 2D position information across the aperture. If strips on one side run perpendicular to those on the other side, then depth-of-interaction information can be obtained. If strips are segmented in either a single-sided or double-sided parallel strip detector then depth-of-interaction information can be obtained and readout rates can be improved.

In the case of double-sided parallel strip detectors (in which opposing strips are parallel) or crossed strip or 2D pixelated array detectors, both electrons and holes produced by a radiation event can be collected to provide 1D positional information between the anode and the cathode sides of the aperture. This 1D positional information is used to determine electronically the sub-aperture spatial resolution.

Sub-aperture spatial resolution can be achieved by measuring either the transit times of electrons and holes to the anodes and cathodes, respectively, or the ratio of anode and cathode signals. A significant benefit may be gained by implementing sub-aperture resolution (e.g., resulting from electronically-defined detector elements) because the edge-on detector aperture height no longer limits spatial resolution along that direction. Furthermore this 1D positional information may, in some situations, be used to estimate meaningful corrections to the expected signal losses as a function of interaction height and thus improve energy resolution. Other benefits include an increase in available image detector volume due to a decrease in the number of edge-on detector physical boundaries (detector material properties typically degrade near the perimeter) and the number of gaps that may be present between edge-on detector planes.

The benefits of sub-aperture resolution (increased spatial resolution, signal loss compensation, fewer readout detectors, increased detector volume) that are possible with edge-on semiconductor detectors can also be attained using scintillator arrays in an edge-on detector geometry. Depth-of-interaction and interaction height information (e.g., for sub-aperture resolution) can be acquired using 1D and 2D scintillator arrays, for example by adding dual-readout (photodetector readout) capability.

The semiconductor detector DOI accuracy is affected by parameters such as the detector depth, electron and hole mobility, signal diffusion, and the number of defects (such as traps) in the bulk semiconductor material. The specific parameters that affect scintillator detector DOI accuracy vary with the DOI measurement technique.

Coupling a 2D photodetector readout array to the side of an edge-on scintillator array permits an analysis of the relative signal strength measured at both ends of individual scintillator elements in the array. By calibrating the relative signal strength versus interaction location in the direction of the aperture (interaction height), sub-aperture resolution can be achieved. With sufficiently fast readout detectors, time-of-flight measurements could also be used to determine the interaction location. Thus, sub-aperture resolution can be attained for 1D and 2D edge-on scintillator detectors, and a 2D, edge-on scintillator array detector can function as a 3D, edge-on scintillator array detector.

In many instances the one-side readout implementations of edge-on SAR designs emulate the face-on DOI designs. In both face-on DOI and edge-on SAR scintillator detector designs, a one-sided or a multi-sided readout can be implemented. Thus, encoding techniques developed for one-sided or two-sided (or multi-sided), face-on DOI scintillator elements can be applied to edge-on SAR scintillator elements. Furthermore, edge-on or edge-on with face-on 2D photodetectors coupled to two or more adjacent faces of a scintillator block (e.g., a block geometry) can be employed to implement a 3D scintillator block detector using encoding techniques (e.g., providing SAR and DOI information).

A problem with face-on DOI scintillator detectors is that Compton scattering of incident radiation is biased in the forward direction such that the probability of detecting the scatter event downstream from the initial off-axis event within the same scintillator may not be small (resulting in an inaccurate DOI estimate). The edge-on SAR scintillator detector design reduces the likelihood that a Compton scatter photon will be detected in the same scintillator for a relatively large range of incident angles. This simplifies the tracking of most subsequent interactions or events after a primary interaction.

The number of edge-on scintillator or semiconductor detector planes required to assemble an edge-on detector module can be reduced by implementing the techniques developed for measuring the depth of interaction (DOI) within face-on scintillator and semiconductor detectors. The benefits of this approach can be illustrated by considering a scenario in which radiation is incident face-on upon the anode or cathode side of a planar semiconductor detector of known depth or thickness (height). The DOI spatial resolution can be determined by measuring either the transit times of electrons and holes to anodes and cathodes, respectively, or the ratio of anode and cathode signals.

Radiation incident approximately perpendicular to the plane or irradiation from the left or right side (approximately parallel to the plane) of an edge-on detector array is also allowed. The side-irradiation geometry may be useful for specific applications. For example, it may be desirable to collimate the radiation so that the detector region near the base and relevant readout electronics are removed from direct irradiation. In addition, irradiation from the right or left side would allow two edge-on detector arrays to be oriented such that one array faces the other array in close proximity. In general, spatial and energy resolution may be enhanced if sub-aperture height information is acquired for edge-on detectors that are irradiated from the side.

FIG. 7 illustrates a perspective view of a detector imaging system 1000 with a one-dimensional (1D) edge-on structured mold detector 700. As shown in FIG. 7, radiation 109 is incident onto the top surface of detector 700, in an edge-on 1D pixelated structured mold detector (silicon block) configuration. Holes 710 of structured mold detector 700 are filled with semiconductor quantum dots or semiconductor detector materials. Channels 720 are filled with attenuation material. These features are not to scale.

In this view, anode face 730 is oriented toward the front of detector 700, showing three separate anode elements 735 separated by the attenuating material in channels 720. Cathode face 740 is oriented toward the back of detector 700, with one or more cathode elements 745. Holes 710 and channels 720 can be etched or micromachined into (e.g., silicon block) detector 700. For example, channels 720 may extend from anode face 730 partially through detector 700 toward cathode face 740, as shown in FIG. 7, or channels 720 may extend completely through detector 700 to (or through) cathode face 740. Alternatively, the front (anode) and back (cathode) faces 730 and 740 can be reversed, and the placement, arrangement, and configurations of channels 720 and holes 710 may vary.

In the particular embodiment of FIG. 7, structured mold detector 700 has holes 710 and channels 720 in which holes 710 are filled with quantum dot materials or other semiconductor detector materials. Channels 720 are filled with an attenuating material to isolate or help isolate neighboring pixels 760. Detector 700 is irradiated in an edge-on configuration, with radiation 109 incident from the top, and front and back faces 730 and 740 of detector 700 are represented by conductive anode and cathode elements 735 and 745, respectively. The top (front-end) and bottom (back-end) layers of detector 700 can also be interchanged, without loss of generality.

For illustrative purposes, an implementation of detector 700 with only a single layer of pixels 760 is shown. A cathode sheet covers back (cathode) face 740 of detector 700, and an anode sheet covers front (anode) face 730. Both the anode sheet on front face 730 and the cathode sheet on back face 740 can be segmented to create individual pixels 760 in detector 700.

Selection of the radius (or other dimensions) of holes 710 and the depth and width of channels 720 is influenced by transport properties of the information carriers (or attenuation properties of the material if used for isolation purposes), and these quantities are selected in a suitable range for detector 700 to identify interactions of radiation 109 within pixels 760. In PbS quantum dots, for example, excitons have a range of about 20 nm, and holes 710 may be selected with a radius of about 50 nm, so that the excitons have acceptable probability of reaching the PbS-silicon heterojunction, for example as described in Urdaneta et al., Quantum dot composite radiation detectors, IEEE Nuclear Science Symposium (2010), which is incorporated by reference herein. Cost is also a design issue for choice of quantum dot or semiconductor detectors materials (such as amorphous or polycrystalline semiconductor materials).

FIG. 8 illustrates a perspective view of a detector imaging system 1000 with a two-dimensional (2D), layered, edge-on structured mold detector (silicon block) 700. In this configuration, channels 720A are provided with relatively low-Z detector material, e.g., in the first or top layer 701 of detector 700, and channels 720B are filled or provided with relatively moderate or high-Z detector material, e.g., in the second or lower layer 702 of detector 700. These features are not to scale.

Channel 720C is provided with a filter material, e.g., between top (low-Z) layer 701 and bottom (moderate or high-Z) layer 702 of detector 700. Anode elements 735 are segmented both by layer 710, 702 and by pixel 706 within each layer 701, 702. Cathode elements 745 can be similarly divided.

In the particular embodiment of FIG. 8, structured mold detector 700 has channels 720A and 720B filled with semiconductor quantum dots or semiconductor detector materials. Hole structures are not necessarily required. As shown, two layers 701, 702 of pixels 706 are provided, with relatively low-Z (or lower-Z) detector materials in channels 720A of first layer 701 (e.g., the top layer, where radiation 109 is incident onto edge-on detector 700), and moderate or relatively high-Z (or higher-Z) materials in second layer 702 (e.g., the bottom layer of detector 700, reached by radiation 109 passing though top layer 701).

For illustrative purposes, the implementations of anode face or layer 730 and cathode face or layer 740 are represented with only one level of pixels 706 per detector layer 701, 702. Alternatively, there may be multiple pixels 706 per detector layer, and/or multiple anode and cathode elements 735 and 745. The selection of channel width (or hole radius) for detector system 700 is influenced by the transport properties of the information carriers, as described above.

FIG. 9 illustrates a perspective view of a multilayer detector system 800 for a CT and/or PET detector imaging system 1000. In this particular example, detector system 800 includes N=4 (four) individual layers: a top or front-end layer 801, two middle layers 802A and 802B, and a bottom or back-end layer 803.

As shown in FIG. 9, the first (top or front-end) layer 801 is formed with an array of edge-on, 2D pixelated detectors, e.g., using a relatively low-Z semiconductor detector material such as silicon. The two middle layers 802A and 802B are formed with arrays of face-on 1D or 2D pixelated detectors, e.g., using a moderate-Z semiconductor detector material such as CZT or CdTe. The bottom (or back-end) layer 803 is formed with an array of edge-on, 2D pixelated detectors, e.g., using a moderate or relatively high-Z scintillator or semiconductor detector material.

The dimensions of pixels within a layer or within different layers may be different. Therefore, the spatial resolution (as well as temporal and energy resolution) properties of individual detector layers (as well as the pixels within a detector layer) need not be the same and are dictated by the imaging requirements as well as cost. This principle applies for one detector layer, two detector layers, three detector layers, four detector layers, etc. For example, although detector layer 803 pixels are depicted with the same surface area (e.g., 1×1) as pixels in detector layers 801 and 802, imaging and cost requirements may indicate (or dictate) a different relationship A×B where each of A and B can be less than, equal to or greater than one (including block detectors).

Depending on application, first (top) layer 801 can be used for CT and PET imaging (e.g., employing a combination of x-ray and Compton scatter interactions). Middle layers 802A and 802B could also be used for CT and/or PET imaging, while bottom layer 803 can be used primarily for PET imaging. In some embodiments, middle layers 802A and 802B are provided as removable/insertable units, which are configured for insertion into and removal from imaging system 1000 between top and bottom layers 801 and 803 of detector 800. Bottom (PET) layer 803 can also be provided in the form of a 3D sub-aperture resolution (SAR) detector, for example as described in Nelson, U.S. Pat. No. 7,635,848. In general, the N detector layer design is appropriate for at least one of Compton (including gamma camera/SPECT) imaging, Compton-PET imaging, PET (including TOF PET) imaging, high resolution CT imaging, CT imaging. Either energy integration or PC or PCE capability can be implemented in individual detector layers depending on the detector capabilities and imaging requirements.

FIG. 10A illustrates a perspective view of an alternate multilayer detector system 800 for, in one implementation, a CT and/or PET detector imaging system 1000. In this particular example, detector system 800 includes N=3 (three) layers: first (top) layer 801, second (middle) layer 802 and third (bottom) layer 803.

As shown in FIG. 10A, first (top or front-end) layer 801 is formed with an array of edge-on, 2D pixelated semiconductor detectors, e.g., using a relatively low-Z semiconductor detector material such as silicon. Second (middle) layer 802 is formed with an array of edge-on pixelated detectors, e.g., using a moderate-Z semiconductor or scintillator detector material (alternatively, layer 802 can be implemented using an array of face-on pixelated detectors). Third (bottom or back-end) layer 803 is formed with an array of edge-on, 2D pixelated detectors, e.g., using a moderate or high-Z detector material. Second (middle) detector layer 802 can also be configured in removable/replaceable form (e.g., if detector layer 803 is not implemented then N=2 layers), and third (bottom) layer can be replaced (for example) by a 3D DOI detector or a 3D SAR detector, as described above for the four-layer embodiment of FIG. 9. Depending upon embodiment, one goal is to provide a three-layer configuration including an edge-on Si, face-on scintillator or semiconductor detector layer and an edge-on semiconductor or scintillator detector layer (e.g., in the front-end and/or back-end layers). A face-on middle detector layer could include an integrator to handle high fluence, e.g., in an x-ray CT imaging system (or combined CT/nuclear medicine system), with the awareness that 511 keV photons may not be entirely contained. Thus, such a face-on layer may provide limited energy resolution, or incorporate photon counting or integration capability, as described herein.

FIG. 10B illustrates a perspective view of CT and/or PET detector imaging system 1000 with a face-on back-end detector layer 803. As shown in FIG. 10B, first (top or front-end) layer 801 and one or more middle layers 802 of detector system 800 are formed with arrays of edge-on, 1D or 2D pixelated detectors. Third (bottom) layer 803 is formed with an array of face-on, 1D or 2D pixelated detectors.

FIG. 10C illustrates a perspective view of a multilayer CT and/or PET detector imaging system 1000 with a face-on back-end detector layer. FIGS. 10B and 10C show one or more middle detector layers 802 of detector system 800 are formed with arrays of edge-on, 1D or 2D pixelated detectors.

FIG. 10D illustrates a perspective view of a multilayer CT and/or PET detector imaging system 1000 with a face-on back-end detector layer. FIG. 10D shows middle detector layer 802 of detector system 800 formed with an array of face-on pixelated detectors.

FIGS. 10B, 10C and 10D illustrate perspective views of CT and/or PET detector imaging system 1000 with an edge-on first (top or front-end) detector layer 801, one or more middle detector layers 802 and a face-on back-end detector layer 803. As shown in FIGS. 10B, 10C and 10D, first (top or front-end) detector layer 801 of detector system 800 is formed with arrays of edge-on, 1D or 2D pixelated detectors. FIGS. 10B and 10C show one or more middle detector layers 802 of detector system 800 formed with arrays of edge-on, 1D or 2D pixelated detectors. FIG. 10D shows middle detector layer 802 of detector system 800 formed with an array of face-on pixelated detectors.

FIGS. 10B, 10C and 10D show a third (bottom) detector layer 803 formed with face-on block, 1D, 2D (or 3D if DOI is implemented) pixelated detectors. As described herein, one or more of the layers could combine edge-on and face-on detector elements (varying at least one of spatial resolution, energy resolution, temporal resolution and stopping power within a detector layer). The dimensions of pixels within a layer or within different layers may be different. The spatial resolution (as well as temporal and energy resolution) properties of individual detector layers (as well as the pixels within a detector layer) need not be the same and are indicated (or dictated) by imaging requirements as well as cost.

This principle applies for one detector layer, two detector layers, three detector layers, four detector layers, etc. For example, in one implementation the surface area of a pixel of third detector layer 803 as shown in FIG. 10C is depicted, for illustrative purposes, as being (approximately) the surface area of a 5×5 array of pixels in first detector layer 801 or middle detector layer 802. In this instance the pixel size employed in detector layers 801 and 802 may be appropriate for high resolution CT imaging (with a pixel surface dimension ranging from approximately 0.2-1.0 mm at this time) whereas the pixel size employed in detector layer 803 may be appropriate for at least one of Compton (including gamma camera/SPECT) imaging, Compton-PET imaging, PET (including TOF PET) imaging, CT imaging. Other implementations may result in third detector layer 803 pixels being smaller, the same or larger than pixels in first detector layer 801 pixels or second detector layer 802.

FIG. 10D depicts detector layer 803 with a pixel area which is more block-like. In this instance detector layer 803 might be comprised of one or more blocks used primarily for timing and/or energy determination for scattered PET photons from detector layer 801 or detector layers 801 and 802 (if detector layer 802 is not present then this represents an implementation of an N=2 detector layer imaging system 1000). (Note that various implementations of block detectors are possible including simple 1D block detectors, 2D block detectors (including, but not limited to, 1D arrays or gamma cameras) and 3D block detectors). Features including fast or very fast temporal response, present in one or both layers of the above-described two layer, coincidence (including TOF) PET detector system implementation, can be present in one or more of the detector layers for the three-layer CT and/or PET, Compton-PET, and Compton (including gamma camera) detector imaging systems. Either energy integration or PC or PCE capability can be implemented in individual detector layers depending on detector capabilities and imaging requirements.

Middle detector layer (or layers) 802 may be provided in a removable configuration and bottom (back-end) detector layer 803 may be replaced with an SAR detector, as described above. In addition, the orientation of first (front-end) and third (back-end) layers 801 and 803 can be interchanged with respect to the direction of incident radiation 109, without loss of generality.

FIG. 11 illustrates a perspective view of a detector imaging system 1000 with a focused two-dimensional (2D), layered, edge-on, pixelated structured mold (silicon block) detector 700. As shown in FIG. 11, individual pixels 706 diverge in width along the direction of incident radiation 109, e.g. with respect to the direction of radiation 109 from a diverging source such as an internal radionuclide or a diverging x-ray beam. This provides pixels 706 and detector 700 with a focused structure geometry, as described herein.

Channels 720A are provided with relatively low-Z detector material, e.g., in the first or top layer 701 of focused detector 700, and channels 720B are filled or provided with relatively moderate or high-Z detector material, e.g., in the second or lower layer 702 of focused detector 700. Channel 720C is filled or provided with a filter material, e.g., between top (low-Z) layer 701 and bottom (moderate or high-Z) layer 702 of focused detector 700. Anode elements 735 and cathode elements 745 can be segmented both by layer 710, 702 and by pixel 706 within each layer 701, 702.

In the particular embodiment of FIG. 11, focused, structured mold detector 700 has channels 720A and 720B filled with semiconductor quantum dot materials or semiconductor detector materials. Hole structures are not necessarily required. As shown, two layers 701, 702 of pixels 706 are provided, with relatively low-Z materials in channels 720A of first layer 701, and moderate or relatively high-Z materials in second layer 702. Anode layer 730 and cathode layer 740 may have multiple anode and cathode elements 735 and 745 per detector layer, with multiple pixels in one or both of layers 701 and 702.

These references are expressly incorporated by reference herein:

-   Bornefalk Hans, Danielsson Mats, Svensson Christer, Image Quality in     Photon Counting-Mode Detector Systems, U.S. Publication No.     2010/0215230 (U.S. Pat. No. 8,378,310). -   Danielsson Mats, Karlsson Staffan, Silicon Detector Assembly for     X-ray Imaging, U.S. Publication No. 2010/0204942 (U.S. Pat. No.     8,183,535). -   Da Via C., et al., Dual Readout—Strip/Pixel Systems, Nucl. Instr.     Meth. A594, p. 7 (2008). -   Knoll, G., Radiation Detection and Measurement, 4th edition, Wiley,     pp. 230, 238, 492 (2010). -   Kroeger R., et al., Three-Compton Telescope: Theory, Simulations,     and Performance, IEEE Trans. Nucl. Sci., Vol. 49(4), pp. 1887-1892     (August 2002). -   Nelson R., Barbaric Z., High Efficiency X-Radiation Converters, U.S.     Pat. No. 4,560,882. -   Nelson R., X-ray Detector for Radiographic Imaging, U.S. Pat. No.     4,937,453. -   Nelson R., Method for Manufacturing a High Resolution Structured     X-ray Detector, U.S. Pat. No. 5,258,145. -   Nelson R., Nelson W., Device and System for Improved Imaging in     Nuclear Medicine and Mammography, U.S. Pat. No. 6,583,420. -   Nelson R., Nelson W., Device and System for Enhanced SPECT, PET, and     Compton Scatter Imaging in Nuclear Medicine, U.S. Pat. No.     7,291,841. -   Nelson R., Devices and Systems for Enhanced SPECT, PET, and Compton     Gamma Cameras, U.S. Pat. No. 7,635,848. -   Nelson R., Nelson W., Slit and Slot Scan, SAR, and Compton Devices     and Systems for Radiation Imaging, U.S. Pat. No. 8,017,906. -   Nelson R., Edge-on SAR Scintillator Devices and Systems for Enhanced     SPECT, PET, and Compton Gamma Cameras, U.S. Pat. No. 8,115,174. -   Nelson R., Edge-on SAR Scintillator Devices and Systems for Enhanced     SPECT, PET, and Compton Gamma Cameras, U.S. Pat. No. 8,115,175. -   Nelson R., Edge-on SAR Scintillator Devices and Systems for Enhanced     SPECT, PET, and Compton Gamma Cameras, U.S. Pat. No. 8,183,533. -   Nelson R., Nelson W., High Resolution Imaging System for Digital     Dentistry, U.S. patent application Ser. No. 13/199,612, filed Sep.     6, 2011 (U.S. Publication No. 2012/0181437). -   Nelson R., Nelson W., Enhanced Resolution Imaging Systems for     Digital Radiography, U.S. patent application Ser. No. 13/507,659,     filed Jul. 18, 2012 (U.S. Publication No. 2013/0028379). -   Parker S., et al., 3DX: an X-ray pixel array detector with active     edges, IEEE Trans. Nucl. Sci. 53 1676-1688 (2006). -   Parker S., et al., Increased speed: 3D silicon sensors; Fast current     amplifiers, IEEE Trans. Nucl. Sci. 58, pp. 404-417 (2011). -   Singh, M., An electronically collimated gamma camera for single     photon emission computed tomography. Part I: Theoretical     considerations and design criteria, Medical Physics Vol. 10(4), pp.     421-427 (July/August 1983). -   Singh, M., Doria D., An electronically collimated gamma camera for     single photon emission computed tomography. Part II: Image     reconstruction and preliminary experimental measurements, Medical     Physics Vol. 10(4), pp. 428-435 (July/August 1983). -   Yu, H. and Wang, G., Compressed sensing based interior tomography,     Phys. Med. Biol., Vol. 54(9): pp. 2791-2805 (2009). -   Urdaneta, M., et al., Quantum dot composite radiation detectors,     IEEE Nuclear Science Symposium (2010).

While this invention has been described with reference to exemplary embodiments, it will be understood by those skilled in the art that various changes can be made and equivalents may be substituted without departing from the spirit and scope thereof. Modifications may also be made to adapt the teachings of the invention to particular problems, technologies, materials, applications and materials, without departing from the essential scope thereof. The invention is not limited to the particular examples that are disclosed herein, but encompasses all embodiments falling within the scope of the appended claims.

While the invention is thus susceptible to various modifications and alternative forms, specific examples thereof have been shown by way of example in the drawings and are herein described in detail. It should be understood, however, that the invention is not to be limited to the particular forms or methods disclosed, but to the contrary, the invention is to cover all modifications, equivalents, and alternatives falling within the spirit and scope of the claims. 

1-49. (canceled)
 50. An imaging apparatus comprising: a front-end detector layer comprising a first array of radiation detector modules configured to image a region of interest within an object; a back-end detector layer comprising a second array of radiation detector modules configured to image the region of interest within the object; detector electronics configured for tracking interactions in the first and second arrays of radiation detector modules using temporal and energy resolution techniques; and a communications link to a computer imaging system configured to image the region of interest within the object by processing readout data from the detector electronics; wherein the front-end and back-end detector layers have temporal and spatial resolution adapted for imaging the region of interest by tracking the interactions from at least one of an x-ray and a gamma radiation source.
 51. The imaging apparatus of claim 50, wherein the temporal and spatial resolution of the front-end and back-end detector layers are adapted for operation of the apparatus as at least one of a Compton camera and x-ray CT imaging system, a PET camera and x-ray CT imaging system, a Compton camera and PET camera and x-ray CT imaging system, an x-ray CT imaging system, a PET camera imaging system, a Compton camera imaging system, a Compton camera and PET camera imaging system, and a gamma camera imaging system.
 52. The imaging apparatus of claim 51, wherein the computer imaging system is configured to image the region of interest by tracking the interactions from the x-ray and gamma radiation source in the front-end and back-end detector layers, respectively.
 53. The imaging apparatus of claim 52, wherein the computer imaging system is configured to image the region of interest by tracking the interactions within or between the front-end and back-end detector layers, respectively.
 54. The imaging apparatus of claim 50, wherein adjacent detector modules in the front-end and back-end detector layers are tilted with respect to one another to achieve a focused detector geometry with respect to diverging radiation from the at least one source, with the adjacent detector modules in the back-end detector layer offset to fill gaps between the adjacent modules in the front-end detector layer and define a substantially continuous detector configuration.
 55. The imaging apparatus of claim 54, wherein the detector modules in at least one of the first array and the second array have a non-uniform readout element pitch.
 56. The imaging apparatus of claim 50, wherein the detector modules in the first array have an edge-on geometry in which at least one of x-ray and gamma radiation are incident upon an edge of a generally planar detector area, the edge having a detector thickness defining an entrance aperture for the radiation and the generally planar area defining an attenuation distance therein.
 57. The imaging apparatus of claim 56, wherein the temporal resolution of the front-end detector layer is adapted for sub-aperture resolution, such that the entrance aperture does not limit spatial resolution in the edge-on geometry.
 58. The imaging apparatus of claim 56, wherein: the detector modules in the second array have an edge-on geometry in which the radiation is incident upon an edge of a generally planar detector area, the edge having a detector thickness defining an entrance aperture for the radiation and the generally planar area defining an attenuation distance therein; or the detector modules in the second array have a face-on geometry in which the radiation is incident upon a face of a generally planar detector area having a detector thickness defining an interaction depth therein.
 59. The imaging apparatus of claim 58, wherein the detector modules in the first array are configured to provide 3D spatial resolution in the front-end detector layer and the detector modules in the second array are configured to provide 2D spatial resolution in the back-end detector layer.
 60. The imaging apparatus of claim 58, wherein the temporal resolution of the back-end detector layer is adapted for depth-of-interaction resolution to estimate a depth of the interactions in the generally planar detector area of the detector modules in the second array.
 61. The imaging apparatus of claim 58, wherein the temporal resolution of the back-end detector layer is adapted for sub-aperture resolution within the generally planar detector area of the detector modules in the second array.
 62. The imaging apparatus of claim 50, wherein the detector modules in at least one of the first array and the second array have a non-uniform readout element pitch configured to balance count rate due to beam hardening with penetration depth.
 63. The imaging apparatus of claim 50, wherein the detector modules in at least one of the first array and the second array are pixelated to define pixel columns having non-uniform pixel size, and wherein pixel length within the pixel columns increases with increasing depth.
 64. The imaging apparatus of claim 63, wherein the detector electronics comprise readout ASICs of progressively slower speeds to read out pixels at greater depth.
 65. The imaging apparatus of claim 50, wherein the computer imaging system is configured to synthesize an effective pixel size by combining the readout data of two or more pixels in at least one of the first or second array of detector modules.
 66. The imaging apparatus of claim 50, wherein the first and second arrays of detector modules comprise scintillator detectors with temporal and spatial resolution selected to track at least one of x-ray interactions and gamma ray interactions from the at least one source.
 67. A radiation imaging system comprising: at least first and second layers of detector modules with temporal, spatial and energy resolution adapted to image a region of interest within a body; detector electronics configured for detecting interactions of radiation in the at least first and second layers of detector modules; and a communication link to a computer imaging system configured to image the region of interest by tracking the interactions in the at least first and second layers of detector modules.
 68. The radiation imaging system of claim 67, wherein the detector modules have at least one different temporal, spatial and energy resolution adapted for the computer imaging system to track interactions in the at least first and second layers, respectively.
 69. The radiation imaging system of claim 67, wherein the detector modules of at least one of the layers have different spatial resolution from the detector modules of at least one other of the layers, the different spatial resolution defined by detector blocks in the at least one layer having a different pixel size than adjacent pixels in the at least one other layer.
 70. The radiation imaging system of claim 69, wherein the at least one layer is a back-end detector layer and the detector blocks cover pluralities of the adjacent pixels in the at least one other layer.
 71. The radiation imaging system of claim 69, wherein the detector blocks comprise at least one of semiconductor detectors, scintillator detectors, straw detectors, gas detectors and liquid detectors with an edge-on or face-on orientation, with or without depth of interaction (DOI) resolution and with or without sub-aperture resolution (SAR).
 72. The radiation imaging system of claim 67, wherein the detector modules have at least one of different temporal resolution and different spatial resolution in the first and second layers.
 73. The radiation imaging system of claim 72, further comprising a partial ring or planar front-end CT detector including the first layer of detector modules and a partial ring or planar back-end PET detector including the second layer of detector modules.
 74. The radiation imaging system of claim 73, further comprising additional front-end detector modules configured to improve at least one of detection efficiency and timing resolution of the back-end PET detector.
 75. The radiation imaging system of claim 73, wherein at least one of the front-end CT detector and the back-end PET detector consists of a single layer of the detector modules.
 76. The radiation imaging system of claim 67, wherein the computer imaging system is configured for tracking both x-ray and gamma ray interactions in operation as at least one of a Compton gamma and x-ray CT imaging system, a PET camera and x-ray CT imaging system, a Compton camera and PET camera and x-ray CT imaging system, an x-ray CT imaging system, a PET camera imaging system, a Compton camera imaging system, a Compton camera and a PET camera imaging system, and a gamma camera imaging system.
 77. The radiation imaging system of claim 76, wherein the detector modules in at least the first layer have an edge-on geometry with respect to the interactions and the detector modules in at least the second layer have either an edge-on geometry or a face-on geometry with respect to the interactions.
 78. The radiation imaging system of claim 77, wherein at least one of the first and second layers of detector modules has a focused geometry with respect to the region of interest, with adjacent offset detector modules filing gaps between adjacent modules in the at least one of the first and second layers to define a substantially continuous detector configuration.
 79. The radiation imaging system of claim 67, wherein the radiation comprises at least one of neutron, photon, and charged particle radiation.
 80. The radiation imaging system of claim 67, wherein the radiation comprises x-ray and gamma ray radiation.
 81. A method of operating a radiation imaging system having at least first and second layers of detector modules with at least one of different temporal, spatial and energy resolution adapted to image a region of interest within a body, the method comprising: utilizing detector electronics configured for detecting interactions from at least one radiation source in each of the at least first and second layers of detector modules; tracking the interactions in each of the at least first and second layers of detector modules; communicating readout data from the detector electronics to a computer imaging system; and imaging the region of interest with the computer imaging system, based on the tracking of the interactions from the at least one radiation source with the different resolution in each of the at least first and second layers of detector modules.
 82. The method of claim 81, wherein the different resolution of the at least first and second layers of detector modules is selected for tracking the interactions from both x-rays and gamma rays.
 83. The method of claim 82, further comprising operating the radiation imaging system as at least one of a Compton gamma and x-ray CT imaging system, a PET camera and x-ray CT imaging system, a Compton camera and PET camera and x-ray CT imaging system, an x-ray CT imaging system, a PET camera imaging system, a Compton camera imaging system, a Compton camera and a PET camera imaging system, and a gamma camera imaging system.
 84. The method of claim 83, further comprising operating first layer of detector modules as part of a front-end CT detector and operating the second layer of detector modules as part of back-end PET detector. 